Photodisruptive laser treatement of crystalline lens

ABSTRACT

Apparatus and methods of treating a hard lens region of an eye with a laser where one method includes identifying a boundary of the hard lens region, selecting a laser-parameter to enable a photo disruptive procedure in the hard lens region and to control a spreading of bubbles in the hard lens region, modifying a mechanical property of a posterior portion of the hard lens region in a proximity of the identified boundary by the photo disruptive procedure, and modifying a mechanical property of a portion anterior to the modified posterior portion of the hard lens region by the photo disruptive procedure. The laser bubbles can be applied to form incisions which are non-transverse to an axis of the eye and intersect the lens fibers.

CROSS-REFERENCE TO RELATED APPLICATION

This application is a continuation of, and thus claims benefit of andpriority from, U.S. patent application Ser. No. 12/343,418, filed Dec.23, 2008 and entitled “Photodisruptive Laser Treatment Of TheCrystalline Lens,” which is a continuation-in-part of, and thus claimsbenefit of and priority to, U.S. patent application Ser. No. 12/205,842,filed Sep. 5, 2008 and entitled “Photodisruptive Laser Treatment of theCrystalline Lens,” which claims benefit of and priority to U.S.provisional application Ser. No. 60/970,454, filed Sep. 6, 2007 andentitled “Photodisruptive Laser Treatment of the Crystalline Lens.” Theentire disclosures of the above-referenced applications are herebyincorporated by reference as part of the disclosure of this application.

BACKGROUND

This application relates to laser eye surgery of the crystalline lensusing photodisruption caused by laser pulses.

Surgical procedures for removal of the crystalline lens utilize varioustechniques to break up the lens into small fragments that can be removedfrom the eye through incisions. Some of these procedures use manualinstruments, ultrasound, heated fluids or lasers. One of the significantdrawbacks of these methods is the need to actually enter the eye withprobes in order to accomplish the fragmentation. This typically requiresmaking large incisions on the lens and limits the precision associatedwith such lens fragmentation techniques.

Photodisruptive laser technology can deliver laser pulses into the lensto optically fragment the lens without insertion of a probe and thus ispotentially a less intrusive procedure, offering higher precision andcontrol.

Laser-induced photodisruption has been already used in the past in laserophthalmic surgery. In the target region the laser ionizes a portion ofthe molecules, eventually releasing gases, which, in an expansion phase,disrupt and break up the lens material in the target region. In somecases Nd:YAG lasers have been employed as the laser sources.

Lens fragmentation via laser-induced photodisruption has also beenproposed. For example, L'Esperance in U.S. Pat. No. 4,538,608 disclosedan apparatus for lens tissue destruction which included a viewingsystem, a laser and a means for optical delivery and scanning of thefocal spot of laser pulses. The laser pulses were focused on theanterior plane of the lens and were moved progressively deeper into thelens to achieve cataract material destruction. In U.S. Pat. No.5,246,435, Bille proposed an alternative approach that focused the laserpulses first in a posterior region of the lens and then move the focusin a posterior to anterior direction. In this method the laser reachedthe target regions with less distortion from the already treatedregions, thus affording greater control. However, various technicalproblems remain unresolved.

SUMMARY

Apparatus and methods of treating a hard lens region of an eye with alaser are provided. Implementations of a method of treating acrystalline lens of an eye with a laser include selecting a surgicalregion of the lens, applying laser pulses to form at least one incisionwithin the selected surgical region, wherein an orientation of theincision is one of an orientation intersecting fibers of the lens and anorientation non-transverse to an axis of the eye, and the incisionmodifies a property of the lens.

In some implementations the non-transverse orientation of the incisionis an orientation substantially parallel to the axis of the eye or anorientation making a less than 90 degree angle with the axis of the eye.

In some implementations a spatial extent of the incision along the axisof the eye is longer than the spatial extent transverse to the axis ofthe eye.

In some implementations the spatial extent along the axis of the eye isin the range of 0.5 mm-12 mm and the spatial extent transverse to theaxis of the eye is in the range of 1-500 microns. In someimplementations the axis of the eye is one of a visual axis, an opticaxis, a line of sight and a pupillary axis.

In some implementations the incision cuts the fibers into partsapproximately at the intersection of the incision and the fibers and themodified property of the lens is a weakening of a biomechanical propertyof the lens.

In some implementations the incision cuts the fibers at or near suturesof the fibers.

In some implementations the incision avoids cutting sutures in the lens.

In some implementations the applying laser pulses includes applying thelaser pulses to generate gas bubbles which form the incision, wherein anorientation of the incision is aligned with a preferential direction ofexpansion of the generated gas bubbles.

In some implementations the applying the laser pulses includes movingthe focal point of the applied laser beam along a posterior to anteriordirection within the lens.

In some implementations the incision has one of an extent at least equalto an extent of a nucleus of the lens, an X-Y diameter in excess of 2 mmand a Z extent in excess of 0.5 mm, and an X-Y diameter in excess of 4mm and a Z extent in excess of 1 mm, wherein the X-Y diameter is ameasure of the spatial extent of the entire incision in the directiontransverse to the axis.

In some implementations the method includes forming no more than oneincision and the laser pulses are applied in a continuous manner to formthe incision without repositioning the laser or interrupting theapplication of the laser.

In some implementations the incision has a form aligned with the axis ofthe eye, the form being of at least one of a cylinder, a set ofconcentric cylinders, a set of cylinders connected by one or moreconnecting line, a curved surface, a cone, a spiral, a layered spiralwith smooth lines connecting layers of the spiral and a tilted cylinder.

In some implementations the incision has a form aligned with the axis ofthe eye, the form being at least one of a plane, two or more crossingplanes, a combination of planes and connecting arcs, and a combinationof planes and cylinders.

In some implementations the applying the laser pulses includes formingincisions in a layer-by layer manner.

In some implementations the forming the incisions in a layer-by-layermanner includes applying laser pulses to target locations within aposterior layer of the lens, the target locations belonging to twoincisions or two segments of the same incision and applying laser pulsesto target locations within a layer anterior to the posterior layer, thetarget locations belonging to the same two incisions or to the same twosegments of the same incision.

In some implementations the applying the laser pulses includes applyingthe laser pulses to form a first ring with a first radius in a posteriorlayer of the lens, applying the laser pulses to form a connector linebetween the first and a second ring in the posterior layer, applying thelaser pulses to form the second ring with a second radius in theposterior layer, and repeating multiple times the formation of the firstring, the second ring and the connector line in layers sequentiallyanterior to the posterior layer, wherein the first rings in thesequential layers form a first cylinder, the second rings form a secondcylinder, the cylinders being connected by the connector lines.

In some implementations the connector lines in sequential layers are oneof aligned to form connector planes and not-aligned from layer to layer.

Some implementations include forming a posterior spiral in a posteriorlayer, forming a smooth connector line starting near an end of thespiral in the posterior layer, the connector line smoothly bending andrising to a central region of a layer anterior to the posterior layerand forming an anterior spiral starting at the end of the smoothconnector line in the central region of the anterior layer.

In some implementations the posterior spiral and the anterior spiral areessentially aligned to form a spiral with an extent in the Z direction.

In some implementations the applying the laser pulses includes selectinglaser-parameters sufficient to create bubbles in the lens, butinsufficient to cause harm to a retina of the eye.

In some implementations the applying the laser pulses includes applyingthe laser pulses with laser-parameters insufficient to fragment the lensto a degree suitable for removal, if the incision were transverse to theaxis of the eye.

In some implementations the laser-parameters include a laser pulseenergy in the range of 0.5 microJ to 50 microJ, a duration of a laserpulse in the range of 0.005 picoseconds to 25 picoseconds, a repetitionrate of applying laser pulses in the range of 1 kHz to 10 MHz, and aseparation distance of target regions of laser pulses in the range of 1micron to 100 microns.

In some implementations the applying the laser pulses includes applyingthe laser pulses with varying energy as the incision is formed.

In some implementations the energy is varied during at least one of a Zdirectional scanning and an X-Y directional scanning.

In some implementations the energy is varied in relation to ameasurement of an optical property of an eye tissue.

Some implementations include forming the incision on a layer-by-layerbasis, wherein one or more layers are at least partially formed along acurved focal plane of a laser delivery system.

In some implementations a Z directional scanner is adjusted at a slowerrate than an X-Y directional scanner when forming a layer of one or moreincisions.

Some implementations further include forming a protection layer in aposterior portion of the lens, positioned to block a large portion ofthe laser pulses applied to form the incision.

In some implementations the incision fragments at least a portion of thelens, the method further including removing the fragmented portion ofthe lens.

In some implementations the applying the laser pulses includes applyinglaser pulses with laser parameters which do not cause lasting damage toa retina of the eye, wherein the laser pulses fragment the lens to adegree suitable for removal and the time of the fragmentation is lessthan a minute.

Some implementations include applying laser pulses to form an incisionin a lens of an eye, wherein the laser pulses are applied by a lasersystem which is configured to scan the laser pulses in the entirenucleus of the lens without interrupting the application of the laserpulses.

In some implementations the incision intersects fibers of the lens.

In some implementations at least segments of the incision areessentially non-transverse to an axis of the eye.

In some implementations the incision is one of a cylinder, a set ofconcentric cylinders, a set of concentric cylinders connected byconnecting lines, a cone, crossing planes, crossing planes connected byarcs, a spiral, and a layered spiral with a smooth line connectinglayers of the spiral.

Some implementations include a laser system for fragmenting acrystalline lens of an eye, including a pulsed laser configured togenerate a laser beam of laser pulses and an optical delivery system,wherein the optical delivery system is configured to apply the laserbeam to create an incision in the lens of the eye with a spatial extentalong an axis of the eye longer than 2 mm and a spatial diametertransverse to the axis of the eye larger than 4 mm without interruptingthe application of the laser.

In some implementations the optical delivery system is configured tomove a focal point of the laser in a posterior to anterior direction ofthe lens.

In some implementations the optical delivery system is configured tocontrol the laser to generate a laser beam with laser-parameterssufficient to create photodisruption in a selected lens region andinsufficient to cause damage to a retina of the eye.

In some implementations the optical delivery system is configured tocontrol the pulsed laser to generate laser pulses with laser-parametersan energy in the range of approximately 0.5 microJ to 50 microJ, aseparation of adjacent target areas in the range of approximately 1micron to 100 microns, a duration in the range of approximately 0.005picoseconds to 25 picoseconds and a repetition rate in the range of 1kHz to 10 MHz.

BRIEF DESCRIPTION OF FIGURES

FIG. 1 illustrates an overview of an eye.

FIG. 2 illustrates a structure of a lens of an eye, including a reducedtransparency region.

FIGS. 3A-B illustrate the generation and spreading of bubbles in aphotodisruptive treatment of a lens.

FIG. 4 illustrates the steps of a photodisruptive treatment of a lens.

FIGS. 5A-C illustrate the steps of a photodisruptive procedure.

FIGS. 5D-K illustrate various configurations of incisions.

FIG. 6 illustrates a step of determining a boundary of the hard lensregion.

FIG. 7 shows an example of an imaging-guided laser surgical system inwhich an imaging module is provided to provide imaging of a target tothe laser control.

FIGS. 8-16 show examples of imaging-guided laser surgical systems withvarying degrees of integration of a laser surgical system and an imagingsystem.

FIG. 17 shows an example of a method for performing laser surgery bysuing an imaging-guided laser surgical system.

FIG. 18 shows an example of an image of an eye from an optical coherencetomography (OCT) imaging module.

FIGS. 19A, 19B, 19C and 19D show two examples of calibration samples forcalibrating an imaging-guided laser surgical system.

FIG. 20 shows an example of attaching a calibration sample material to apatent interface in an imaging-guided laser surgical system forcalibrating the system.

FIG. 21 shows an example of reference marks created by a surgical laserbeam on a glass surface.

FIG. 22 shows an example of the calibration process and thepost-calibration surgical operation for an imaging-guided laser surgicalsystem.

FIGS. 23A and 23B show two operation modes of an exemplaryimaging-guided laser surgical system that captures images oflaser-induced photodisruption byproduct and the target issue to guidelaser alignment.

FIGS. 24 and 25 show examples of laser alignment operations inimaging-guided laser surgical systems.

FIG. 26 shows an exemplary laser surgical system based on the laseralignment using the image of the photodisruption byproduct.

DETAILED DESCRIPTION

FIG. 1 illustrates the overall structure of the eye. The incident lightpropagates through the optical path which includes the cornea, theanterior chamber, the pupil, the posterior chamber, the lens and thevitreous humor. These optical elements guide the light on the retina.

FIG. 2 illustrates a lens 100 in more detail. The lens 100 is sometimesreferred to as crystalline lens because of the α, β, and γ crystallineproteins which make up about 90% of the lens. The crystalline lens hasmultiple optical functions in the eye, including its dynamic focusingcapability. The lens is a unique tissue of the human body in that itcontinues to grow in size during gestation, after birth and throughoutlife. The lens grows by developing new lens fiber cells starting fromthe germinal center located on the equatorial periphery of the lens. Thelens fibers are long, thin, transparent cells, with diameters typicallybetween 4-7 microns and lengths of up to 12 mm. The oldest lens fibersare located centrally within the lens, forming the nucleus. The nucleus101 can be further subdivided into embryonic, fetal and adult nuclearzones. The new growth around the nucleus 101, referred to as cortex 103,develops in concentric ellipsoid layers, regions, or zones. Because thenucleus 101 and the cortex 103 are formed at different stages of thehuman development, their optical properties are distinct. While the lensincreases in diameter over time, it may also undergo compaction so thatthe properties of the nucleus 101 and the surrounding cortex 103 maybecome even more different (Freel et al, BMC Ophthalmology 2003, vol. 3,p. 1).

As a result of this complex growth process, a typical lens 100 includesa harder nucleus 101 with an axial extent of about 2 mm, surrounded by asofter cortex 103 of axial width of 1-2 mm, contained by a much thinnercapsule membrane 105, of typical width of about 20 microns. These valuesmay change from person to person to a considerable degree.

Lens fiber cells undergo progressive loss of cytoplasmic elements withthe passage of time. Since no blood veins or lymphatics reach the lensto supply its inner zone, with advancing age the optical clarity,flexibility and other functional properties of the lens sometimesdeteriorate.

FIG. 2 illustrates, that in some circumstances, including long-termultraviolet exposure, exposure to radiation in general, denaturation oflens proteins, secondary effects of diseases such as diabetes,hypertension and advanced age, a region of the nucleus 101 can become areduced transparency region 107. The reduced transparency region 107 isusually a centrally located region of the lens (Sweeney et al Exp. Eyeres, 1998, vol. 67, p. 587-95). This progressive loss of transparencyoften correlates with the development of the most common type ofcataract in the same region, as well as with an increase of lensstiffness. This process may occur with advancing age in a gradualfashion from the peripheral to the central portion of the lens (Heys etal Molecular Vision 2004, vol. 10, p. 956-63). One result of suchchanges is the development of presbyopia and cataract that increase inseverity and incidence with age.

The reduced transparency region 107 can be removed via cataract surgery.A common procedure is to make an incision into the capsule of the cloudylens (capsulotomy) and surgically remove the interior, i.e. the cortexand the nucleus, while leaving the lens capsule intact. This is theso-called extra capsular surgery. While the cortex exhibits viscousfluid dynamics and thus can be removed by aspiration or even simplesuction, the nucleus is too hard for this approach and is typicallyremoved as a whole. Finally, a plastic “intraocular” lens is ofteninserted as a replacement into the capsule. This procedure requiresmaking a quite large incision, sometimes up to 12 mm. Creating incisionsof this size can lead to a variety of problems, as described below.

In some methods, the use of ultrasound waves was introduced intocataract surgery. In this “phacoemulsification” procedure one or moresmaller incisions are made on the capsule 105 and an ultrasoundagitator, or “phaco-probe” is introduced into the lens. Operating theagitator or phaco-probe emulsifies the nucleus, which allows the removalof the emulsified nucleus via aspiration through an incision smallerthan the previous technique.

However, even the phacoemulsification technique requires making anincision on the capsule 105, sometimes up to 7 mm. The procedure canleave extensive unintended modifications in its wake: the treated eyecan exhibit extensive stigmatism and a residual or secondary refractiveor other error. This latter often necessitates a follow-up refractive orother surgery or device.

In recent developments, considerable effort was focused on developing alarge variety of the intraocular lenses for insertion into the capsule105. The examples include even bifocal lenses. However, there wasn'tmuch progress in the area of improving the removal process involving thelens 100 or the nucleus 101.

Implementations of the present application include photodisruptivemethods instead of phacoemulsification to break up a hard lens region109. Since no phaco probe is inserted into the lens 100, a much smallerincision is necessitated only for the subsequent aspiration of thebroken-up nucleus. This reduces the unintended secondary effects, andcan reduce the percentage of patients who need secondary refractive orother surgery.

The hard lens region 109 often coincides with the nucleus 101. However,numerous variations may occur. E.g. the outermost soft layers of thenucleus may be removable by aspiration or even suction and thus may notrequire photodisruptive methods. In other cases, only thecataract-impacted portion of the eye may be disrupted for subsequentremoval. In yet other cases it may be desired that only a portion of thenucleus 101 is disrupted, when the nucleus is only sculpted and notremoved. To express the broader scope of the contemplated variations,all these regions will be jointly referred to as the hard lens region109. The nucleus 101 is only one embodiment of the hard lens region 109.

In some cases this hard lens region 109 may occupy an ellipsoid-likeregion of approximately 6-8 mm in equatorial diameter and approximately2-3.5 mm in axial diameter, or extent. The size of this hard lens region109 may be different for different patients, for different diseases andfor different procedures.

In a laser-induced lens fragmentation process, laser pulses ionize aportion of the molecules in the target region. This may lead to anavalanche of secondary ionization processes above a “plasma threshold”.In many surgical procedures a large amount of energy is transferred tothe target region in short bursts. These concentrated energy pulses maygasify the ionized region, leading to the formation of cavitationbubbles. These bubbles may form with a diameter of a few microns andexpand with supersonic speeds to 50-100 microns. As the expansion of thebubbles decelerates to subsonic speeds, they may induce shockwaves inthe surrounding tissue, causing secondary disruption.

Both the bubbles themselves and the induced shockwaves carry out a goalof the procedure: the disruption, fragmentation or emulsification of thetargeted hard lens region 109 without having made an incision on thecapsule 105. The disrupted hard lens region 109 can then be removedthrough a much smaller incision, possibly without inserting a surgicaldevice into the lens itself.

However, the photodisruption decreases the transparency of the affectedregion. Remarkably, the lens of the eye has the highest density ofproteins of all tissues, yet it is transparent. For this same reason,however, the transparency of the lens is particularly sensitive tostructural changes, including the presence of bubbles and damage byshockwaves.

If the application of the laser pulses starts with focusing them in thefrontal or anterior region of the lens and then the focus is moveddeeper towards the posterior region, the cavitation bubbles and theaccompanying reduced transparency tissue can be in the optical path ofthe subsequent laser pulses, blocking, attenuating or scattering them.This may diminish the precision and control of the application of thesubsequent laser pulses, as well as reduce the energy pulse actuallydelivered to the deeper posterior regions of the lens. Therefore, theefficiency of laser-based eye surgical procedures can be enhanced bymethods in which the bubbles generated by the early laser pulses do notblock the optical path of the subsequent laser pulses.

Various approaches, including the technique of U.S. Pat. No. 5,246,435,do not provide an effective way of addressing the above adverseinterference by bubbles produced by preceding laser pulses. Thus, priormethods often require the use of additional lens fragmentationtechniques in addition to the photodisruption by laser.

In recognition of the above technical problem and based on theinvestigation of the distinct properties of the various lens regions andthe laser pulse parameters on the generation and spreading of cavitationbubbles, the techniques, apparatus and systems described in thisapplication can be used to effectively fragment the crystalline lens bylaser pulses with reduced interference from the bubbles induced bypreceding laser pulses. Subsequently, the removal of a portion of or theentirety of the crystalline lens can be achieved via aspiration withreduced or no need of other lens fragmentation or modificationtechniques.

FIG. 3 illustrates that the hard lens region 109 with differenttransport, optical and biomechanical properties has significantimplications for the photodisruptive fragmentation techniques. Onesignificant limitation of the various laser-based lens fragmentationtechniques is the hard-to-control spread of gas bubbles that may occurduring the photodisruption that can reduce the effectiveness of thesubsequent laser pulses to carry out their intended function.

FIG. 3A illustrates that a laser beam 110, which is focused to a smallfocal or target area can generate a small gas bubble 111.

FIG. 3B illustrates that the resistance against the spread of thiscavitation bubble 111 can vary from layer to layer of the lens 100.Inside the nucleus 101, the small bubble 111 may simply expand into abigger bubble 112. It may also generate shockwaves around the bubble, asshown at 114. Moreover, if the expanding bubble reaches thenucleus-cortex boundary, as bubble 116 does, then the gas can expandextensively in the softer cortex region 103. Any of these extendedgaseous bubbles can disturb, absorb, scatter or even block thesubsequent laser pulses, directed to fragment the hard lens region.

In addition, there may be pre-existing channels in the hard lens regionthat may allow the generated gas to move into the softer lens regionsand interfere with further pulse delivery. Such channels may be locatedalong suture lines, where lens fibers meet. Avoidance of these andadjacent areas may also be employed to reduce gas spread. In addition,pulse properties may be modified in these areas to further reduce gasspread. Such areas can be identified preoperatively or alternatively,intra-operative identification of such channels can allow the procedureto be altered.

Methods, which first attempt to remove the softer peripheral layers,including the cortex 103 and attempt to remove the harder nucleus 101afterwards, face considerable drawbacks, because the initial removal ofthe peripheral layers may leave behind a disrupted, unclear opticalpath, making the subsequent fragmentation of the harder nucleus 101 bylasers difficult.

It is noteworthy that laser-disruption techniques developed for otherareas of the eye, such as the cornea, cannot be practiced on the lenswithout substantial modification. One reason for this is that the corneais a highly layered structure, inhibiting the spread and movement ofbubbles very efficiently. Thus, the spread of bubbles posesqualitatively lesser challenges in the cornea than in the softer layersof the lens including the nucleus itself.

The resistance of the various lens regions against the spreading of thegas bubbles 111 depends on numerous individual characteristics of eachpatient including the age of the patient. The spread of gas can also beinfluenced by the particular laser parameters applied to the target.

FIG. 4 illustrates an implementation of a photo-disruptive eye-surgicalprocess 200 developed from the above considerations.

FIGS. 5A-K illustrate various embodiments of the method of FIG. 4.

In step 210 a boundary 252 of the hard lens region 109 may be determinedfrom measuring a mechanical or optical characteristic of the lens 100.Implementations may include this step 210 because if the laser pulsesare applied outside the hard lens region 109, the generated bubbles mayexpand considerably and in a hard-to-control manner. Therefore, someimplementations may include first a determination of the boundary of thehard lens region 109 so that the laser pulses can be focused inside thehard lens region 109.

FIG. 6 shows an implementation of step 210 based on mechanicalcharacteristics of the bubbles. A string of probe-bubbles 290 may begenerated in the lens 100, for example, substantially parallel with amain axis of the eye, separated by a suitable distance, such as 10 to100 microns. Other bubble strings can be generated in other areas of thelens. As shown, since the harder nucleus 101 shows more resistanceagainst the expansion of the probe-bubbles, the probe-bubbles 290-1inside the hard nucleus 101 may expand slower. By the same token, thecortex 103 may exert less resistance against the expansion of thebubbles and thus the probe-bubbles 290-2 outside the nucleus 101, in thecortex 103 may expand faster. A portion of the boundary 252 between thenucleus 101 and the cortex 103 can then be identified as the line orregion separating slow-expanding probe-bubbles 290-1 from fast-expandingprobe-bubbles 290-2.

The expansion of the probe-bubbles 290 and the line separating theslow-expanding probe-bubbles 290-1 from the fast-expanding probe-bubbles290-2 may be observed and tracked by an optical observation method. Manysuch methods are known, including all kinds of imaging techniques.Mapping out or otherwise recording these separation points or lines canbe used to establish the boundary 252 between the softer lens regionsand the hard lens region 109. This implementation of step 210 can bepre-operative, i.e. performed prior to the surgical procedure, orintra-operative, i.e. performed as an early phase of the surgicalprocedure.

Several other methods can be applied for step 210 as well. For example,optical or structural measurements can be performed prior to thesurgical procedure on the patient. Or, some database can be used, whichcorrelates some other measureable characteristic of the eye to the sizeof the nucleus, e.g. using an age-dependent algorithm. In some cases anexplicit calculation can be employed as well. In some cases even datafrom cadavers can be utilized. It is also possible to generate the abovebubble string, and then apply an ultrasound agitation, and observe theinduced oscillation of the bubbles, especially their frequency. Fromthese observations, the hardness of the surrounding tissue can beinferred as well.

In some cases the method of Optical Coherence Tomography (OCT) can beutilized in step 210. Among other aspects, OCT can measure the opacityof the imaged tissue. From this measurement, the size of the bubbles andthe hardness of the region can be inferred once again.

Finally, the hard lens region 109 can be selected based on some otherconsideration, e.g. when only the cataract region is to be removed, orthe nucleus is to be sculpted only. All of these methods are within thescope of step 210 of FIG. 4, and are illustrated in FIG. 5A with thedotted line indicating the boundary 252 of the hard lens region 109.

FIG. 4 illustrates that step 220 may include selecting a laser parameterbetween a disruption-threshold and a spread-threshold. The laserparameters of the laser pulses 110 can be selected to be above thedisruption-threshold for generating the photodisruption in the hard lensregion 109. The laser parameters can be selected to be below thespread-threshold that creates uncontrolled spreading of the gas producedby the photodisruption.

These disruption- and spread-thresholds can be demonstrated e.g. in thecase of the spatial separation between two adjacent target points of thelaser pulses. If the generated bubbles are closer than a lowerspread-threshold distance, then the bubbles may coalesce, forming abigger bubble. These larger bubbles are likely to expand faster and in aharder-to-control manner. On the other hand, if the bubbles are fartherthan the upper disruption-threshold, then they may not achieve theintended photodisruption or fragmentation of the target tissue. In somecases the range of bubble separation between these thresholds can bebetween 1 micron and 50 microns.

The duration of the laser pulses may also have analogous disruption- andspread-thresholds. In some implementations the duration may vary in therange of 0.01 picoseconds to 50 picoseconds. In some patients particularresults were achieved in the pulse duration range of 100 femtoseconds to2 picoseconds. In some implementations, the laser energy per pulse canvary between the thresholds of 1 microJ and 25 microJ. The laser pulserepetition rate can vary between the thresholds of 10 kHz and 100 MHz.

The energy, target separation, duration and repeat frequency of thelaser pulses can also be selected based on a preoperative measurement oflens optical or structural properties. Alternatively, the selection ofthe laser energy and the target separation can be based on apreoperative measurement of the overall lens dimensions and the use ofan age-dependant algorithm, calculations, cadaver measurements, ordatabases.

FIG. 4 illustrates that in step 230 a mechanical property of a posteriorportion of the hard lens region can be modified in the proximity of theidentified boundary 252 by a photodisruptive procedure.

FIG. 5B illustrates an embodiment of step 230, where a set of bubbles isgenerated by initial laser pulses 110-1 in a posterior portion 254 ofthe hard lens region 109, in the proximity of the boundary 252. Themodifying the mechanical property may include that the generated bubblesphotodisrupt, fragment, or even emulsify the tissue of the posteriorportion 254 of the nucleus 101, thus modifying some of its mechanicalproperties.

FIG. 4 illustrates that in step 240 a mechanical property of a portionanterior to the already modified posterior portion can be modified by aphotodisruptive procedure.

FIG. 5C illustrates an embodiment of step 240, where a second set ofbubbles are generated by subsequent laser pulses 110-2 in a region 256which is anterior to the already modified region 254.

In implementations of the method these photodisruptive steps 240 can berepeatedly applied by moving the focal or target region of the laserbeam 110 along a direction from the posterior of the hard lens region109 to the anterior of the hard lens region 109. This sequence of thephotodisruptive steps 240 controls and limits the buildup and spread ofbubbles in the optical path of the subsequent laser pulses 110-2. Theseimplementations allow the subsequent laser pulses 110-2 to deliveressentially their entire energy to the target area, allow for bettercontrol of the subsequent pulses, as well as clearer imaging of thesurgical area for the benefit of the person conducting the procedure.

Steps 210-240 may be followed by the removal of the fragmented,disrupted, emulsified or otherwise modified hard lens regions 109, ifrequired or desired. One method of removing the fragmented, disrupted,or otherwise modified regions is to create one or more small openings,or incisions in the lens capsule 105, and then insert an aspirationprobe to remove the fragmented material. In other implementations,simple suction can extract the fragmented material, as well as thenon-fragmented viscous material, such as the cortex 103, withoutinserting a probe into the capsule.

When laser pulses are applied to the hard lens region 109 from theposterior to anterior direction, between the disruption- and thespread-thresholds, they can optically modify, photodisrupt, or fragmentthe structure of the treated hard lens region 109 to facilitate lensmaterial removal while also reducing the spread of gas and bubblesduring placement of these initial and subsequent laser pulses. Thecharacteristics of the hard lens region 109 can vary from patient topatient though, thus the disruption-threshold and spread-threshold laserparameters may need to be determined from patient to patient.

In some implementations, the energy of the laser beam can be adjusted asthe focal point is moved in the posterior-to-anterior direction. Toreach the anterior layers, the laser beam passes through less materialand thus a laser beam with less energy can achieve the same disruptionin the target tissue. Accordingly, applying a laser beam with a constantenergy may generate an increasing amount of gas as the laser is moved inthe anterior direction. To avoid the generation and subsequent spread ofsuch an excess amount of gas, in some implementations the laser energycan be reduced as the laser is moved in the posterior-to-anteriordirection. In other implementations, the applied laser energy can alsobe adjusted as the laser is scanning in the X-Y transverse direction, asthe amount of material the laser passes through also varies as thescanning proceeds in the X-Y transverse direction.

In some implementations, the rate of reduction of the applied energy canbe calculated from an imaging procedure, which is sensitive e.g. to anoptical density or a scattering of the imaged target tissue.

Additional laser pulses can be applied subsequent to the initial laserapplication, at target positions in the lens outside the initiallytreated zone in the central region of the lens. The gas and bubblescreated by these subsequent laser pulses can either permeate in thetreated central region of the lens without uncontrollably spreading inthe lens, or can spread into the lens tissue outside the initiallytreated zone. As such, the gas produced by photodisruption in theperipheral areas of the lens does not block effective treatment of thehard lens region 109. The laser treated hard lens region and theperipheral lens material which may or may not be treated with the laserdepending on need can be removed from the eye via aspiration, with orwithout additional lens tissue breakup using mechanical, suction,ultrasonic, laser, heated fluid or other means. In anotherimplementation, only the treated region is removed via aspiration, withor without additional lens tissue breakup using mechanical, suction,ultrasonic, laser, heated fluid or other means.

FIGS. 5D-K illustrate other implementations of the eye surgical method200. To set the stage for the description of these methods, a note onterminology. In the following the terminology “an axis of the eye” willbe used extensively. There are several ways to define an axis of theeye. The axes of the eye can be categorized e.g. according to the GrandY. L. Physiological Optics (Springer-Verlag, New York, 1980) as follows:

Optical axis: Line passing through the optical center of the cornea andthe lens;

Visual axis: Line passing from the point of fixation to the image on thecenter of the retina called fovea;

Line of Sight: Line passing from the object point through the center ofthe entrance of the pupil; and

Pupillary axis: Line passing perpendicularly through the center of thecornea and the center of the entrance of the pupil.

In practice these axes are often quite close to each other. Further,compromise axes can be defined as well, e.g. an axis which lies betweenany two or three of the above axes. In the rest of this disclosure thescope of the term “the axis of the eye” will include any one of thesedefinitions. The axis of the eye will be also referred to as the Z axis.In typical implementations the laser beam can also be oriented along theZ axis. However, other implementations where the laser beam makes anangle with the Z axis are also within the scope of the described method.The two directions transverse to the Z axis will be sometimes termed Xand Y axes, following customary terminology.

General aspects of these implementations include the following.

First, these implementations benefit from the recognition that a primarysource of the biomechanical strength of the lens is based on its fibers.As described above, the fibers are an elongated, hardened, essentiallytransparent tissue within the lens, which grow around the center of theeye in a somewhat irregular manner, typically starting from theequatorial plane. The length of fibers can vary widely. In some casesthe length falls within the range of 1-10 mm. However, fiber lengthsoutside this range can also occur. The fibers can be joined at sutures.In various contexts the fiber structure of the lens has been describedas layered, as an onion structure and as a ball of yarn. Close to theaxis of the eye, the fiber layers are typically oriented in a mannernear perpendicular, or transverse, to the axis.

The fiber-rich central region forms the nucleus. Accordingly, to aconsiderable degree the biomechanical strength of the nucleus isprovided by the fibers and their layers, which are nearperpendicular/transverse to the axis of the eye.

Second, as described below in more detail, there is an improvedunderstanding of the dynamics and expansion of the laser generatedcavitation bubbles which make up the incision, indicating that theyexpand quite differently parallel and transverse to the fiber layers inthe lens. Implementations of the surgical methods exploit thesedifferences to improve the efficiency and control of the surgicalprocess.

Third, these implementations also benefit from the availability of newand improved eye-surgical laser systems, which are capable of scanning alarge fraction of the surgical area, in some cases the entire area,without repositioning. As described below, this feature may offersubstantial positive aspects.

In the light of the above described three developments, someimplementations of the eye surgical method differ from existing methodsat least in the following aspects:

(i) The incisions are non-transverse: Close to the center of the eyeincisions maybe positioned and oriented in directions which arenon-transverse to the axis of the eye. Accordingly, the extent of theincisions can be long along the Z axis and smaller in the X-Y plane.

In some embodiments, the incisions can be essentially parallel to theaxis of the eye. Examples include cylinders, whose axis is essentiallyparallel to the axis of the eye. In some cases the length of thecylinder can be between 0.5 mm to 12 mm in the Z direction and theextent in the X-Y plane, essentially the thickness of the incision, canbe in the range of 0.1-500 microns.

A shared characteristic of some of these embodiments is that theindividual incisions or features have a longer spatial extent in the Zdirection, or axis-parallel direction, than in the X-Y direction, ortransverse direction. In the case of e.g. cylindrical incisions (seebelow), the length of the cylinder along the Z axis is longer than thethickness of its wall in the X-Y direction. The term “extent in the X-Ydirection” will be used to refer to that of the single incision itself,such as its thickness, and not an overall dimension of the geometricform of the incision, e.g. the diameter of a cylinder. In someembodiments, the spatial extent of the incisions in the Z direction canbe in the range of 0.5-10 mm, the extent in the X-Y direction, i.e. theX-Y thickness can be in the range of 1-500 microns, and the X-Y diameterof the incision can be in the range of 2-10 mm. The spatial extent ofthe individual incisions can be chosen depending on the number ofparallel incisions and their separation.

Other embodiments can be practiced as well, where the incisions makesome angle with the axis of the eye, e.g. in the form of a cone, or atilted cylinder, or any other form, non-transverse to the axis of theeye. Non-transverse incisions with piece-wise transverse sections arealso within the scope of these implementations.

(ii) The incisions cut fibers: Close to the axis of the eye, because ofthe non-transverse orientation of the incisions, the incisions cutthrough some of the fibers of the lens, as the fibers and their layersare typically close to transverse to the axis of the eye. In peripheralareas of the lens the fibers and their layers tilt/bend away from thetransverse direction. Accordingly, in these peripheral regions theincisions themselves can be oriented in a direction which still cuts thenon-transverse fibers. Since the fibers are a primary source of thebiomechanical strength of the lens, cutting through the fibers reducesthe biomechanical strength of the lens effectively.

(iii) The orientation of the incision offers superior gas management:The impact of the laser beam creates miniscule bubbles in the targettissue. Experiments reveal that these bubbles undergo a two stageexpansion. During an initial fast expansion, the bubbles may expand atsupersonic speeds, and thus can be very efficient atfragmenting/disrupting the surrounding tissue. This fast expansion istypically anisotropic and occurs mostly in the direction of the laserbeam, i.e. approximately the Z direction. The second stage of theexpansion is slower, and typically occurs towards the softer tissue,i.e. between the fiber layers, in the transverse direction. During thisslow transverse expansion, bubbles often coalesce into bigger bubbles,which can obscure the optical path of subsequent laser pulses,considerably undermining the control and efficiency of the procedure.

In existing methods, which create transverse incisions, the fast,Z-directional bubble expansion does not help creating the transverseincision, and therefore the surgeon has to create the bubbles much moreclosely to each other.

In contrast, in implementations of the present method, creatingincisions approximately in the Z direction, the anisotropy of the fastbubble expansion is put to good use, as it allows the surgeon to createfewer bubbles spaced farther apart in the Z direction, since the bubbleswill fast expand in the Z direction and fragment the tissue betweenneighboring bubbles efficiently.

Such a reduction of the necessary number of bubbles or the equivalentreduction of laser energy in the present method is a criticaldifference, as most of the laser beam, after having left the lens,reaches the retina. The retina, being a photosensitive tissue, maysuffer substantial damage because of the impact of this laser beam. Toachieve a fast and substantial fragmentation of the lens tissue, theenergy of the laser is often chosen to be close to values which candamage the retina. Therefore, the reduction of the necessary number ofbubbles or the energy per pulse of the laser in the present method canmean the difference between damaging the retina and leaving it intact.

Furthermore, the present method also offers advantages regarding thesecond, slower bubble expansion. During this stage the bubbles expand inthe transverse direction. As described above, these bubbles, especiallywhen coalescing together, can substantially and disadvantageouslyobscure the target area, reducing the efficiency and control of thesurgical procedure.

In the present method, the surgeon can create the Z-directed incisionslayer-by-layer (see FIG. 5F), creating only lines of bubbles in eachlayer. Therefore, the surgeon can move the focus of the laser fasterthan the transverse expansion of the previously created bubbles.

In contrast, existing methods create transverse incisions, i.e. thesurgeon has to create bubbles covering entire areas, returningrepeatedly to regions which have been passed earlier. In these methodsit is hard or near impossible for the surgeon to move the laser fasterthan the expanding bubbles, or to avoid returning to previously impactedareas. In fact the surgeon is regularly forced to operate in the areaobscured by the expanding bubbles, leading to a considerable reductionof precision and control over the surgical procedure.

(iv) The incisions avoid sutures in some implementations: As mentionedbefore, fibers typically come together, or end, in sutures. Thesesutures often form planar structures, parallel to the Z axis. It hasbeen observed that in some cases bubbles expand particularly fast alongsutures. Such a too-fast expansion may result in obscuring or cloudingthe optical path even if Z directional incisions are formed, thuspossibly reducing control and precision. Therefore, some implementationsof the method create incisions away from sutures.

At the same time, other implementations may be based on the observationthat the sutures provide a structural framework for the fibers, and thuscutting through the sutures may be particularly effective in reducingthe biomechanical stability of the lens. This benefit has to be weighedagainst the above mentioned drawback of fast-expanding bubbles along thesutures. Depending on the comparative cost-benefit analysis and theother requirements of the method, some implementations may avoid makingincisions at or near the sutures, while others may cut through some ofthe sutures.

(v) Making fewer incisions applies less energy to the eye: Since thefiber-cutting incisions are quite efficient in reducing thebiomechanical strength of the lens, a reduced number of incisions arecapable of achieving the extent of tissue fragmentation necessary forthe objectives of the eye surgery. Reduced number of incisions can beapplied in shorter time, thereby applying less energy to the eye.Therefore, these surgical methods deposit a reduced amount of energy inthe eye, thus e.g. reducing the potential risk to light sensitivetissue, such as the retina by this method.

In some implementations, an eye surgical method making transverseincisions may require 150-160 seconds to achieve the fragmenting of thelens to the degree which is reached in only 45-50 seconds with methodswhich make essentially axis-parallel incisions.

This factor of 3-4 reduction of surgery time can be quite beneficial,since often surgical patients develop hard-to-control eye movementsafter about 120 seconds, necessitating the abandonment of the surgicalprocedure. The just-described reduction of surgery time can mean thedifference between the successful completion of the surgery and itsabandonment.

Equivalently, this time reduction can be converted into reducing theenergy deposited by the laser by a factor of 3-4 in fiber-cuttingmethods during comparable surgery times, thereby substantially reducingthe potential for damage in the retina.

(vi) Incisions are few and extended: The eye surgical method can beperformed with surgical instruments which are configured to createincisions with unprecedented spatial extent. In some implementations ofthe surgical instrument this extent can be 0.5-10 mm in the Z direction,in some cases 2-4 mm, and 2-8 mm in the X-Y plane. This large spatialextent of the incisions imparts several positive features to thesurgical method, as described below.

(vii) Extended incisions have fewer acceleration/deceleration regions:When making an individual incision, at the beginning of the incision themovement of the laser-focal-point typically accelerates from zero to theregular scanning speed. While the laser is accelerating, it may depositenergy at a higher rate or higher density to the eye, possibly leadingto damage in the light sensitive tissues, such as the retina. The sameapplies at the end of incisions, when the laser-focal-point isdecelerating, again possibly damaging the retina. Therefore, methodswhich utilize longer incisions reduce the number ofacceleration/deceleration regions, thus reducing the potential fordamage to the light sensitive tissues in these regions in contrast tomethods which use a large number of minute incisions.

Existing surgical systems are unable to avoid this problem, as theirscanning range in the Z and X-Y directions is considerably less than theentire surgical region. In some existing systems the X-Y scanning rangecan be 1-2 mm and the Z scanning range can be 0.5 mm, which issubstantially less than the entire surgical region of the lens, such asthe size of the nucleus. Typically, the nucleus has a Z extent of 2-4 mmand an X-Y diameter of 6-10 mm. This limitation of the existing systemsrequires that the surgeon make a large number of smaller incisions, withlots of acceleration/deceleration regions. Once the laser scannerreaches its maximum range when making an incision, the surgeon has tostop the scanning via a deceleration, then reposition the laser scannerpointing to a new scan-start point and start a new incision with anacceleration region. Thus methods using existing laser surgical systemsinvolve creating a number of acceleration/deceleration regions, with theconcomitant problems.

In contrast, implementations of the present method may benefit from theavailability of improved laser systems, which can have a considerablyextended X-Y scanning range of 2-10 mm and Z range of 0.5-10 mm.Therefore, implementations of the present method may involve making onlya few incisions, thus generating only few of the problematicacceleration/deceleration regions.

In particular, some surgical laser systems may be capable of scanningthe entire surgical region. With such systems, the lens-surgery mayinvolve creating only one, uninterrupted extended incision, thus havingthe lowest possible number of acceleration/deceleration regions.

Here it is mentioned that surgical laser systems with larger X-Yscanning ranges have been described before. However, these systems wereused for surgery on the cornea. There are crucial differences betweenlens surgery and cornea surgery, as during lens surgery both the imaginglight off the target and the applied laser propagate through opticallyactive regions: the cornea, the antechamber and part of the lens itself.Propagation through these regions deflects the light substantially bothbecause their differing index of refraction as well as their varyingcurvature. Therefore, considerable corrections and calculations arerequired by the surgical equipment and its operator to point the laserto its indented target region.

Further, the laser beam needs to be not only pointed but also focused onthe target. A convergent beam is, by definition, extended off its focuspoint, or target. Therefore, prior to reaching the target, differentsections of the converging laser beam propagate through regions of theeye with different optical properties and different curvatures, posing asecond level of challenges.

In contrast, the cornea is the outermost optically active layer of theeye. Therefore, neither the pointing nor the focusing of the laser beamposes a hard challenge. Further, the problems arising from the curvatureof the cornea can be minimized e.g. by applanating it, i.e. making thecornea essentially flat by applying various contact lenses and devices.In contrast, applanating the lens is quite challenging and presently noproposal is available how to achieve this.

Because of all the described hard challenges, corneal surgical lasersystems are qualitatively simpler than lens surgical lasers. This iswell supported by the fact that even though corneal surgical systemswere suggested about 40 years ago, none have been adapted successfullyfor lens surgery to date.

(viii) Extended incisions pose less stringent requirements forsynchronization: Surgical lasers typically have a beam controller,configured to switch the laser beam on and off, or control the laser viaa shutter mechanism. This beam controller is synchronized with the beamscanner as the laser is switched off when the beam scanner reaches itsmaximum range, or the end of the incision intended by the surgeon. Thesesurgeries require synchronization between the beam scanner, the beamcontroller and the surgeon's actions. In surgical methods which employ alarge number of small incisions, this need for synchronization posesstringent requirements on the beam controller and beam scanner. Incontrast, surgical methods which employ few and extended incisionsimpose considerably less stringent synchronization requirements.

(ix) Extended incisions have fewer transient laser fronts: When thelaser is switched on to start a new incision, the initial front of thelaser may have transients which are less-well-controlled. These laserfronts may carry a less-well-controlled amount of energy and may be lesswell focused on the intended target region. Surgical methods usinglonger incisions and thus employing fewer switch on/off events reducethe number of such less-well-controlled laser fronts and transients,increasing the control over the tissue fragmentation.

(x) Extended incisions minimize z-scanner movement: Minimizing speed andthe acceleration of the scanner mechanism along the Z axis isparticularly important because the limits of speed and accelerationalong the Z axis are more stringent than along the X and Y axes. Whilescanning in the transverse X-Y direction is achieved in some embodimentsby rotating small and light scanning mirrors, Z axis scanningcustomarily involves translating a lens or a lens group of the deliverysystem linearly along the optical axis. This lens or lens group isusually heavier than the scanning mirrors and thus has a higher inertia.Therefore, moving this lens or lens group fast can be more difficultthan moving the X-Y scanning mirrors. Extended incisions place lessdemanding requirements on the movement of the Z-scanner.

Aspects (vi) to (x) highlight that eye surgical systems which arecapable of scanning the laser beam in a more extended range and thus arecapable of making longer incisions without repositioning, offersubstantial positive aspects over systems which are capable of shorterincisions only.

In particular, laser scanners which can scan the laser beam across theentire surgical region without interruption or repositioning can avoidmost of the deficiencies of prior systems, which require suchrepositioning, as described in points (vi)-(x).

In some implementations the laser pulses can be applied withlaser-parameters which are sufficient to create bubbles in the lens, butinsufficient to cause harm to a retina of the eye.

Because of the enhanced efficiency of the incisions in the abovedescribed surgical method to weaken the biomechanical properties of thelens, in some implementations the laser pulses can be applied withlaser-parameters which would have been insufficient to fragment the lensto a degree suitable for removal, had the pulses been used to form anincision transverse to the axis of the eye.

Laser parameters in various implementations may fall into this“insufficient-to-fragment” range if a laser pulse energy is in the rangeof 0.5 microJ to 50 microJ, a duration of a laser pulse is in the rangeof 0.005 picoseconds to 25 picoseconds, a repetition rate of applyinglaser pulses is in the range of 1 kHz to 10 MHz, and a separationdistance of target regions of laser pulses is in the range of 1 micronto 100 microns.

For all the above reasons, laser-formed incisions which are dominantlynon-transverse to the visual axis and thus cut through layers of lensfibers, weaken the biomechanical strength of the lens qualitatively moreefficiently than incisions which are transverse to the axis and thus cutonly few fibers or none at all. Therefore, implementations of thismethod require considerably less power, shorter application time orlower repetition rate for the surgical laser pulses. Due to thisefficiency of these implementations, the treatment times to fragment thelens can be reduced by a factor of 3-4 or more. Further, implementationsbenefit in a multiplicity of ways from new and improved surgical systemswhich allow the scanning of the entire surgical region withoutinterruption or repositioning.

FIGS. 5D-K illustrate various implementations of the surgical method.The surgical method may start with the surgeon selecting a surgicalregion of the eye to be treated. Next, the surgeon may design theprocedure by selecting the location of the non-transverse incisions tobe made. Then, the surgeon can form non-transverse incisions in thesurgical region by the fast and repeated application of laser pulses.During the application of the laser pulses, the focus of the laserpulses can be moved in a posterior to anterior direction so that thepreviously formed bubbles do not obscure the target region thesubsequent laser pulses are to be applied.

FIGS. 5D-K illustrate various incisions in the lens 100, created byvarious implementation of the surgical method.

FIG. 5D illustrates two views of dominantly transverse incisions, formedby a large number of bubbles generated in transverse layers. Theseincisions will also be referred to as transverse incisions. The viewshown on the left side of FIG. 5D illustrates the layers of bubbleswhich make up the transverse layers 260-i from the side, highlightingthe X-Z plane of the lens. The view shown on the right side of FIG. 5Dillustrates the same from the top, highlighting the X-Y plane.

FIG. 5E illustrates two views of a dominantly axis-parallel orZ-directional cylindrical incision. The view shown on the left side ofFIG. 5E illustrates the bubbles which make up concentric axis-parallelcylinders 262-i from the side, highlighting the X-Z plane of the lens.The bubbles are shown only on the outermost cylinder for clarity. Theview shown on the right side of FIG. 5E illustrates the same from thetop, highlighting the X-Y plane. While in typical implementations thebubbles are densely packed, the figures show the bubbles only sparselyand on selected cylinders to avoid clutter. As described above,analogous implementations can utilize any related geometrical form whichis non-transverse to the optical axis, including incisions of the formof a cone, tilted cylinder, bulging or bending shape. These incisionstypically cut through the fibers of the lens.

FIG. 5F illustrates steps of generating several cylindrical incisions.This particular implementation involves three cylinders, but others mayinvolve any number of cylinders which are capable of achieving thesurgical goal, e.g. the photodisruption of the lens.

In some implementations the cylinders are formed layer-by layersimultaneously, i.e. in parallel. These implementations face less of aproblem regarding the subsequent laser targeting being hindered by theexpansion of the earlier formed bubbles.

To start with, the surgeon may decide the posterior-most depth of theincisions. A guiding principle may be to make sure that the incision issafely within the lens, and therefore the capsule is not accidentallypierced by the method, leading to undesirable consequences.

Then the surgeon may apply laser pulses to form a ring of bubbles with adiameter of e.g. the outermost cylinder 262-1, to form theposterior-most ring of cylinder 262-1. When the laser focal point ismoved along the entire ring and arrives back to the starting point SP,the surgeon may move the focus of the laser along connector-line 263-1towards the center until it reaches the next cylinder 262-2. The focusof the laser is then moved again to form the posterior-most ring ofcylinder 262-2. Finally, again using the connector-line 263-1, theposterior-most ring of the innermost cylinder 262-3 is formed the sameway.

An aspect of this method is that all these steps were carried out bycontinuously applying the laser, in effect creating one incision.Therefore, at no time is the surgeon forced to switch off the laserbeam, thus avoiding the problems described in points (vi)-(x) above. Inother implementations more than one incision is made, but still only afew of them, and not a large number of minute incisions.

Next, the surgeon can move the focus of the laser in aposterior-to-anterior direction, and start forming the second layer ofrings of the three cylinders 262-1, . . . 262-3. Thus, layer-by-layer,the three cylinders 262-1, . . . 262-3 can be formed essentiallysimultaneously.

In the implementation of FIG. 5F, the connector-lines 263 are aligned indifferent ring-layers.

FIG. 5F illustrates a different implementation, where the connectorlines are not aligned in different layers. Visibly, the connector lines263-1, . . . , 263-3 in ring layers 1, 2, 3 can be rotated relative toeach other. For clarity the connector lines in the lower layers areshown with dotted lines.

These implementations simplify the scanning pattern of the pulses andavoid the need for special measures to block or turn off the laser whilemoving from one incision to another. In such cases, the effectiveness ofthe fragmentation by the incisions may be further enhanced by thealternation of the position and/or orientation of the connectingsegments.

FIG. 5G illustrates two views of a Cross Plane embodiment. The viewshown on the left side of FIG. 5G illustrates one of the two crossplanes 265-1 from the side, highlighting the X-Z plane of the lens. Thecentral column of bubbles shown with bold lines indicates the other ofthe two cross planes 265-2, pointing out from the X-Z plane.

The view shown on the right side of FIG. 5G illustrates the same twocross planes 265-1 and 265-2 from the top, highlighting the X-Y plane.

In practice, these two cross planes again can be formed by alayer-by-layer approach, i.e. forming the posterior-most row of bubblesof cross plane 265-1, then move the focal point of the laser along anarc to the starting point of the posterior-most row of the other crossplane 265-2 and form that row. As the cross-planes are being formedlayer-by-layer, the arcs can form a cylinder around the cross planes. Inthis sense, this implementation creates an integratedcross-plane/cylinder structure.

A large number of variations and combinations of the aboveimplementations are possible.

FIG. 5H illustrates that e.g. instead of two cross planes 3, 4, 6, etccross planes 265-i can be created, forming “slices” or “wedges” of thecylindrical surgical region.

FIG. 5I illustrates two spiral shaped incisions. A spiral shapedincision 267, where no large angle redirection is involved in theformation of the incision, is shown on the left side of FIG. I.

The right side of FIG. I illustrates a multi-layer spiral incision. Inthis implementation when a spiral incision 267-1 is completed in afirst, posterior layer, the surgeon can move the focal point of thesurgical laser to the central starting point of the spiral in a second,anterior layer following a smooth and gently rising connecting line 268,and then start creating the spiral 267-2 in this second anterior layer.This smooth connecting line 268, indicated by the solid dots, can be anapproximate semi-circle, or any one of a large number of similarlysmooth curves. Such smooth connecting lines reduce the acceleration ofthe focal point, providing for a more even application of laser energyinto the target tissue.

FIG. 5J illustrates that the focal plane 271 is typically curved inoptical systems unless (any suitable portion of) the optics 273 of thelaser delivery system is corrected for field curvature. In mostuncorrected optical systems the curvature is positive, i.e. the focallength is longer for axial beams 275-1 and shorter for off-axis beams275-2, as shown in FIG. 5J.

When the intended incision is a straight transverse line or and extendedtransverse planar cut, the servo motor driving the Z-scanner (the “Zservo”) can be continuously adjusted in order to compensate thedistorting effect of field curvature. However, since the transverse X-Yscanning speed may be much higher than the Z scanning speed because ofthe higher inertia associated with the Z scanning, the Z servo may notbe able to adjust the focus of the laser beam in the Z direction at thehigh speed of the X-Y scanner.

FIG. 5J illustrates an implementation which does not require adjustingthe Z-scanner at the X-Y transverse scanning rate. In thisimplementation an incision 276 is formed which follows the curvature ofthe focal plane 271 of the laser delivery optics 273. The incision 276can be any of the previously described non-transverse lines,non-transverse planar cuts, layers of spirals, nested cylinders orcrossed planes. When any of these implementations are formed on alayer-by-layer basis, the incisions in several or all of the layers mayfollow the curvature of the focal plane 271, thus reducing oreliminating the need to move the Z servo at the rate of the X-Y scanner.Therefore, these implementations can be operated at the fast X-Ytransverse scanning speed instead of the slower Z-scanning speed.

In yet other various embodiments the incisions can take a wide varietyof shapes including straight planes, curved planes, cones, tiltedcylinders, any type of shapes which are not transverse to the z axis,incisions which have portions which are transverse to the z axis,various crossing patterns and any combination of these patterns. Suchshapes can be connected by interconnecting planes that further fragmentthe lens tissue, while also potentially easing the delivery of laserpulses by reducing the need to shutter the laser or make large movementswith the scanning system.

After cutting the fibers by the laser-formed incisions, the cut fiberscan be removed with a variety of techniques, including hydro-dissection,manual fragmentation, the application of ultrasound, aspiration, or acombination of these or other methods.

FIG. 5K illustrates yet another composite implementation. In thisimplementation, a shield layer 280 can be implemented in theposterior-most region of the lens, to a substantial degree transverse tothe axis of the eye. One of the functions of this protection layer 280is to protect the retina from the negative effects of the laserirradiation used for forming the incisions 262-i.

FIGS. 7-26 illustrate embodiments of a laser surgery system in relationto the above photodisruptive laser treatment.

One important aspect of laser surgical procedures is precise control andaiming of a laser beam, e.g., the beam position and beam focusing. Lasersurgery systems can be designed to include laser control and aimingtools to precisely target laser pulses to a particular target inside thetissue. In various nanosecond photodisruptive laser surgical systems,such as the Nd:YAG laser systems, the required level of targetingprecision is relatively low. This is in part because the laser energyused is relatively high and thus the affected tissue area is alsorelatively large, often covering an impacted area with a dimension inthe hundreds of microns. The time between laser pulses in such systemstend to be long and manual controlled targeting is feasible and iscommonly used. One example of such manual targeting mechanisms is abiomicroscope to visualize the target tissue in combination with asecondary laser source used as an aiming beam. The surgeon manuallymoves the focus of a laser focusing lens, usually with a joystickcontrol, which is parfocal (with or without an offset) with their imagethrough the microscope, so that the surgical beam or aiming beam is inbest focus on the intended target.

Such techniques designed for use with low repetition rate laser surgicalsystems may be difficult to use with high repetition rate lasersoperating at thousands of shots per second and relatively low energy perpulse. In surgical operations with high repetition rate lasers, muchhigher precision may be required due to the small effects of each singlelaser pulse and much higher positioning speed may be required due to theneed to deliver thousands of pulses to new treatment areas very quickly.

Examples of high repetition rate pulsed lasers for laser surgicalsystems include pulsed lasers at a pulse repetition rate of thousands ofshots per second or higher with relatively low energy per pulse. Suchlasers use relatively low energy per pulse to localize the tissue effectcaused by laser-induced photodisruption, e.g., the impacted tissue areaby photodisruption on the order of microns or tens of microns. Thislocalized tissue effect can improve the precision of the laser surgeryand can be desirable in certain surgical procedures such as laser eyesurgery. In one example of such surgery, placement of many hundred,thousands or millions of contiguous, nearly contiguous or pulsesseparated by known distances, can be used to achieve certain desiredsurgical effects, such as tissue incisions, separations orfragmentation.

Various surgical procedures using high repetition rate photodisruptivelaser surgical systems with shorter laser pulse durations may requirehigh precision in positioning each pulse in the target tissue undersurgery both in an absolute position with respect to a target locationon the target tissue and a relative position with respect to precedingpulses. For example, in some cases, laser pulses may be required to bedelivered next to each other with an accuracy of a few microns withinthe time between pulses, which can be on the order of microseconds.Because the time between two sequential pulses is short and theprecision requirement for the pulse alignment is high, manual targetingas used in low repetition rate pulsed laser systems may be no longeradequate or feasible.

One technique to facilitate and control precise, high speed positioningrequirement for delivery of laser pulses into the tissue is attaching aapplanation plate made of a transparent material such as a glass with apredefined contact surface to the tissue so that the contact surface ofthe applanation plate forms a well-defined optical interface with thetissue. This well-defined interface can facilitate transmission andfocusing of laser light into the tissue to control or reduce opticalaberrations or variations (such as due to specific eye opticalproperties or changes that occur with surface drying) that are mostcritical at the air-tissue interface, which in the eye is at theanterior surface of the cornea. Contact lenses can be designed forvarious applications and targets inside the eye and other tissues,including ones that are disposable or reusable. The contact glass orapplanation plate on the surface of the target tissue can be used as areference plate relative to which laser pulses are focused through theadjustment of focusing elements within the laser delivery system. Thisuse of a contact glass or applanation plate provides better control ofthe optical qualities of the tissue surface and thus allow laser pulsesto be accurately placed at a high speed at a desired location(interaction point) in the target tissue relative to the applanationreference plate with little optical distortion of the laser pulses.

One way for implementing an applanation plate on an eye is to use theapplanation plate to provide a positional reference for delivering thelaser pulses into a target tissue in the eye. This use of theapplanation plate as a positional reference can be based on the knowndesired location of laser pulse focus in the target with sufficientaccuracy prior to firing the laser pulses and that the relativepositions of the reference plate and the individual internal tissuetarget must remain constant during laser firing. In addition, thismethod can require the focusing of the laser pulse to the desiredlocation to be predictable and repeatable between eyes or in differentregions within the same eye. In practical systems, it can be difficultto use the applanation plate as a positional reference to preciselylocalize laser pulses intraocularly because the above conditions may notbe met in practical systems.

For example, if the crystalline lens is the surgical target, the precisedistance from the reference plate on the surface of the eye to thetarget tends to vary due to the presence of collapsible structures, suchas the cornea itself, the anterior chamber, and the iris. Not only istheir considerable variability in the distance between the applanatedcornea and the lens between individual eyes, but there can also bevariation within the same eye depending on the specific surgical andapplanation technique used by the surgeon. In addition, there can bemovement of the targeted lens tissue relative to the applanated surfaceduring the firing of the thousands of laser pulses required forachieving the surgical effect, further complicating the accuratedelivery of pulses. In addition, structure within the eye may move dueto the build-up of photodisruptive byproducts, such as cavitationbubbles. For example, laser pulses delivered to the crystalline lens cancause the lens capsule to bulge forward, requiring adjustment to targetthis tissue for subsequent placement of laser pulses. Furthermore, itcan be difficult to use computer models and simulations to predict, withsufficient accuracy, the actual location of target tissues after theapplanation plate is removed and to adjust placement of laser pulses toachieve the desired localization without applanation in part because ofthe highly variable nature of applanation effects, which can depend onfactors particular to the individual cornea or eye, and the specificsurgical and applanation technique used by a surgeon.

In addition to the physical effects of applanation thatdisproportionably affect the localization of internal tissue structures,in some surgical processes, it may be desirable for a targeting systemto anticipate or account for nonlinear characteristics ofphotodisruption which can occur when using short pulse duration lasers.Photodisruption is a nonlinear optical process in the tissue materialand can cause complications in beam alignment and beam targeting. Forexample, one of the nonlinear optical effects in the tissue materialwhen interacting with laser pulses during the photodisruption is thatthe refractive index of the tissue material experienced by the laserpulses is no longer a constant but varies with the intensity of thelight. Because the intensity of the light in the laser pulses variesspatially within the pulsed laser beam, along and across the propagationdirection of the pulsed laser beam, the refractive index of the tissuematerial also varies spatially. One consequence of this nonlinearrefractive index is self-focusing or self-defocusing in the tissuematerial that changes the actual focus of and shifts the position of thefocus of the pulsed laser beam inside the tissue. Therefore, a precisealignment of the pulsed laser beam to each target tissue position in thetarget tissue may also need to account for the nonlinear optical effectsof the tissue material on the laser beam. In addition, it may benecessary to adjust the energy in each pulse to deliver the samephysical effect in different regions of the target due to differentphysical characteristics, such as hardness, or due to opticalconsiderations such as absorption or scattering of laser pulse lighttraveling to a particular region. In such cases, the differences innon-linear focusing effects between pulses of different energy valuescan also affect the laser alignment and laser targeting of the surgicalpulses.

Thus, in surgical procedures in which non superficial structures aretargeted, the use of a superficial applanation plate based on apositional reference provided by the applanation plate may beinsufficient to achieve precise laser pulse localization in internaltissue targets. The use of the applanation plate as the reference forguiding laser delivery may require measurements of the thickness andplate position of the applanation plate with high accuracy because thedeviation from nominal is directly translated into a depth precisionerror. High precision applanation lenses can be costly, especially forsingle use disposable applanation plates.

The techniques, apparatus and systems described in this document can beimplemented in ways that provide a targeting mechanism to deliver shortlaser pulses through an applanation plate to a desired localizationinside the eye with precision and at a high speed without requiring theknown desired location of laser pulse focus in the target withsufficient accuracy prior to firing the laser pulses and withoutrequiring that the relative positions of the reference plate and theindividual internal tissue target remain constant during laser firing.As such, the present techniques, apparatus and systems can be used forvarious surgical procedures where physical conditions of the targettissue under surgery tend to vary and are difficult to control and thedimension of the applanation lens tends to vary from one lens toanother. The present techniques, apparatus and systems may also be usedfor other surgical targets where distortion or movement of the surgicaltarget relative to the surface of the structure is present or non-linearoptical effects make precise targeting problematic. Examples for suchsurgical targets different from the eye include the heart, deeper tissuein the skin and others.

The present techniques, apparatus and systems can be implemented in waysthat maintain the benefits provided by an applanation plate, including,for example, control of the surface shape and hydration, as well asreductions in optical distortion, while providing for the preciselocalization of photodisruption to internal structures of the applanatedsurface. This can be accomplished through the use of an integratedimaging device to localize the target tissue relative to the focusingoptics of the delivery system. The exact type of imaging device andmethod can vary and may depend on the specific nature of the target andthe required level of precision.

An applanation lens may be implemented with another mechanism to fix theeye to prevent translational and rotational movement of the eye.Examples of such fixation devices include the use of a suction ring.Such fixation mechanism can also lead to unwanted distortion or movementof the surgical target. The present techniques, apparatus and systemscan be implemented to provide, for high repetition rate laser surgicalsystems that utilize an applanation plate and/or fixation means fornon-superficial surgical targets, a targeting mechanism to provideintraoperative imaging to monitor such distortion and movement of thesurgical target.

Specific examples of laser surgical techniques, apparatus and systemsare described below to use an optical imaging module to capture imagesof a target tissue to obtain positioning information of the targettissue, e.g., before and during a surgical procedure. Such obtainedpositioning information can be used to control the positioning andfocusing of the surgical laser beam in the target tissue to provideaccurate control of the placement of the surgical laser pulses in highrepetition rate laser systems. In one implementation, during a surgicalprocedure, the images obtained by the optical imaging module can be usedto dynamically control the position and focus of the surgical laserbeam. In addition, lower energy and shot laser pulses tend to besensitive to optical distortions, such a laser surgical system canimplement an applanation plate with a flat or curved interface attachingto the target tissue to provide a controlled and stable opticalinterface between the target tissue and the surgical laser system and tomitigate and control optical aberrations at the tissue surface.

As an example, FIG. 7 shows a laser surgical system based on opticalimaging and applanation. This system includes a pulsed laser 1010 toproduce a surgical laser beam 1012 of laser pulses, and an optics module1020 to receive the surgical laser beam 1012 and to focus and direct thefocused surgical laser beam 1022 onto a target tissue 1001, such as aneye, to cause photodisruption in the target tissue 1001. An applanationplate can be provided to be in contact with the target tissue 1001 toproduce an interface for transmitting laser pulses to the target tissue1001 and light coming from the target tissue 1001 through the interface.Notably, an optical imaging device 1030 is provided to capture light1050 carrying target tissue images 1050 or imaging information from thetarget tissue 1001 to create an image of the target tissue 1001. Theimaging signal 1032 from the imaging device 1030 is sent to a systemcontrol module 1040. The system control module 1040 operates to processthe captured images from the image device 1030 and to control the opticsmodule 1020 to adjust the position and focus of the surgical laser beam1022 at the target tissue 101 based on information from the capturedimages. The optics module 120 can include one or more lenses and mayfurther include one or more reflectors. A control actuator can beincluded in the optics module 1020 to adjust the focusing and the beamdirection in response to a beam control signal 1044 from the systemcontrol module 1040. The control module 1040 can also control the pulsedlaser 1010 via a laser control signal 1042.

The optical imaging device 1030 may be implemented to produce an opticalimaging beam that is separate from the surgical laser beam 1022 to probethe target tissue 1001 and the returned light of the optical imagingbeam is captured by the optical imaging device 1030 to obtain the imagesof the target tissue 1001. One example of such an optical imaging device1030 is an optical coherence tomography (OCT) imaging module which usestwo imaging beams, one probe beam directed to the target tissue 1001thought the applanation plate and another reference beam in a referenceoptical path, to optically interfere with each other to obtain images ofthe target tissue 1001. In other implementations, the optical imagingdevice 1030 can use scattered or reflected light from the target tissue1001 to capture images without sending a designated optical imaging beamto the target tissue 1001. For example, the imaging device 1030 can be asensing array of sensing elements such as CCD or CMS sensors. Forexample, the images of photodisruption byproduct produced by thesurgical laser beam 1022 may be captured by the optical imaging device1030 for controlling the focusing and positioning of the surgical laserbeam 1022. When the optical imaging device 1030 is designed to guidesurgical laser beam alignment using the image of the photodisruptionbyproduct, the optical imaging device 1030 captures images of thephotodisruption byproduct such as the laser-induced bubbles or cavities.The imaging device 1030 may also be an ultrasound imaging device tocapture images based on acoustic images.

The system control module 1040 processes image data from the imagingdevice 1030 that includes the position offset information for thephotodisruption byproduct from the target tissue position in the targettissue 1001. Based on the information obtained from the image, the beamcontrol signal 1044 is generated to control the optics module 1020 whichadjusts the laser beam 1022. A digital processing unit can be includedin the system control module 1040 to perform various data processing forthe laser alignment.

The above techniques and systems can be used deliver high repetitionrate laser pulses to subsurface targets with a precision required forcontiguous pulse placement, as needed for cutting or volume disruptionapplications. This can be accomplished with or without the use of areference source on the surface of the target and can take into accountmovement of the target following applanation or during placement oflaser pulses.

The applanation plate in the present systems is provided to facilitateand control precise, high speed positioning requirement for delivery oflaser pulses into the tissue. Such an applanation plate can be made of atransparent material such as a glass with a predefined contact surfaceto the tissue so that the contact surface of the applanation plate formsa well-defined optical interface with the tissue. This well-definedinterface can facilitate transmission and focusing of laser light intothe tissue to control or reduce optical aberrations or variations (suchas due to specific eye optical properties or changes that occur withsurface drying) that are most critical at the air-tissue interface,which in the eye is at the anterior surface of the cornea. A number ofcontact lenses have been designed for various applications and targetsinside the eye and other tissues, including ones that are disposable orreusable. The contact glass or applanation plate on the surface of thetarget tissue is used as a reference plate relative to which laserpulses are focused through the adjustment of focusing elements withinthe laser delivery system relative. Inherent in such an approach are theadditional benefits afforded by the contact glass or applanation platedescribed previously, including control of the optical qualities of thetissue surface. Accordingly, laser pulses can be accurately placed at ahigh speed at a desired location (interaction point) in the targettissue relative to the applanation reference plate with little opticaldistortion of the laser pulses.

The optical imaging device 1030 in FIG. 7 captures images of the targettissue 1001 via the applanation plate. The control module 1040 processesthe captured images to extract position information from the capturedimages and uses the extracted position information as a positionreference or guide to control the position and focus of the surgicallaser beam 1022. This imaging-guided laser surgery can be implementedwithout relying on the applanation plate as a position reference becausethe position of the applanation plate tends to change due to variousfactors as discussed above. Hence, although the applanation plateprovides a desired optical interface for the surgical laser beam toenter the target tissue and to capture images of the target tissue, itmay be difficult to use the applanation plate as a position reference toalign and control the position and focus of the surgical laser beam foraccurate delivery of laser pulses. The imaging-guided control of theposition and focus of the surgical laser beam based on the imagingdevice 1030 and the control module 1040 allows the images of the targettissue 1001, e.g., images of inner structures of an eye, to be used asposition references, without using the applanation plate to provide aposition reference.

In addition to the physical effects of applanation thatdisproportionably affect the localization of internal tissue structures,in some surgical processes, it may be desirable for a targeting systemto anticipate or account for nonlinear characteristics ofphotodisruption which can occur when using short pulse duration lasers.Photodisruption can cause complications in beam alignment and beamtargeting. For example, one of the nonlinear optical effects in thetissue material when interacting with laser pulses during thephotodisruption is that the refractive index of the tissue materialexperienced by the laser pulses is no longer a constant but varies withthe intensity of the light. Because the intensity of the light in thelaser pulses varies spatially within the pulsed laser beam, along andacross the propagation direction of the pulsed laser beam, therefractive index of the tissue material also varies spatially. Oneconsequence of this nonlinear refractive index is self-focusing orself-defocusing in the tissue material that changes the actual focus ofand shifts the position of the focus of the pulsed laser beam inside thetissue. Therefore, a precise alignment of the pulsed laser beam to eachtarget tissue position in the target tissue may also need to account forthe nonlinear optical effects of the tissue material on the laser beam.The energy of the laser pulses may be adjusted to deliver the samephysical effect in different regions of the target due to differentphysical characteristics, such as hardness, or due to opticalconsiderations such as absorption or scattering of laser pulse lighttraveling to a particular region. In such cases, the differences innon-linear focusing effects between pulses of different energy valuescan also affect the laser alignment and laser targeting of the surgicalpulses. In this regard, the direct images obtained from the target issueby the imaging device 1030 can be used to monitor the actual position ofthe surgical laser beam 1022 which reflects the combined effects ofnonlinear optical effects in the target tissue and provide positionreferences for control of the beam position and beam focus.

The techniques, apparatus and systems described here can be used incombination of an applanation plate to provide control of the surfaceshape and hydration, to reduce optical distortion, and provide forprecise localization of photodisruption to internal structures throughthe applanated surface. The imaging-guided control of the beam positionand focus described here can be applied to surgical systems andprocedures that use means other than applanation plates to fix the eye,including the use of a suction ring which can lead to distortion ormovement of the surgical target.

The following sections first describe examples of techniques, apparatusand systems for automated imaging-guided laser surgery based on varyingdegrees of integration of imaging functions into the laser control partof the systems. An optical or other modality imaging module, such as anOCT imaging module, can be used to direct a probe light or other type ofbeam to capture images of a target tissue, e.g., structures inside aneye. A surgical laser beam of laser pulses such as femtosecond orpicosecond laser pulses can be guided by position information in thecaptured images to control the focusing and positioning of the surgicallaser beam during the surgery. Both the surgical laser beam and theprobe light beam can be sequentially or simultaneously directed to thetarget tissue during the surgery so that the surgical laser beam can becontrolled based on the captured images to ensure precision and accuracyof the surgery.

Such imaging-guided laser surgery can be used to provide accurate andprecise focusing and positioning of the surgical laser beam during thesurgery because the beam control is based on images of the target tissuefollowing applanation or fixation of the target tissue, either justbefore or nearly simultaneously with delivery of the surgical pulses.Notably, certain parameters of the target tissue such as the eyemeasured before the surgery may change during the surgery due to variousfactor such as preparation of the target tissue (e.g., fixating the eyeto an applanation lens) and the alternation of the target tissue by thesurgical operations. Therefore, measured parameters of the target tissueprior to such factors and/or the surgery may no longer reflect thephysical conditions of the target tissue during the surgery. The presentimaging-guided laser surgery can mitigate technical issues in connectionwith such changes for focusing and positioning the surgical laser beambefore and during the surgery.

The present imaging-guided laser surgery may be effectively used foraccurate surgical operations inside a target tissue. For example, whenperforming laser surgery inside the eye, laser light is focused insidethe eye to achieve optical breakdown of the targeted tissue and suchoptical interactions can change the internal structure of the eye. Forexample, the crystalline lens can change its position, shape, thicknessand diameter during accommodation, not only between prior measurementand surgery but also during surgery. Attaching the eye to the surgicalinstrument by mechanical means can change the shape of the eye in a notwell defined way and further, the change can vary during surgery due tovarious factors, e.g., patient movement. Attaching means includefixating the eye with a suction ring and applanating the eye with a flator curved lens. These changes amount to as much as a few millimeters.Mechanically referencing and fixating the surface of the eye such as theanterior surface of the cornea or limbus does not work well whenperforming precision laser microsurgery inside the eye.

The post preparation or near simultaneous imaging in the presentimaging-guided laser surgery can be used to establish three-dimensionalpositional references between the inside features of the eye and thesurgical instrument in an environment where changes occur prior to andduring surgery. The positional reference information provided by theimaging prior to applanation and/or fixation of the eye, or during theactual surgery reflects the effects of changes in the eye and thusprovides an accurate guidance to focusing and positioning of thesurgical laser beam. A system based on the present imaging-guided lasersurgery can be configured to be simple in structure and cost efficient.For example, a portion of the optical components associated with guidingthe surgical laser beam can be shared with optical components forguiding the probe light beam for imaging the target tissue to simplifythe device structure and the optical alignment and calibration of theimaging and surgical light beams.

The imaging-guided laser surgical systems described below use the OCTimaging as an example of an imaging instrument and other non-OCT imagingdevices may also be used to capture images for controlling the surgicallasers during the surgery. As illustrated in the examples below,integration of the imaging and surgical subsystems can be implemented tovarious degrees. In the simplest form without integrating hardware, theimaging and laser surgical subsystems are separated and can communicateto one another through interfaces. Such designs can provide flexibilityin the designs of the two subsystems. Integration between the twosubsystems, by some hardware components such as a patient interface,further expands the functionality by offering better registration ofsurgical area to the hardware components, more accurate calibration andmay improve workflow. As the degree of integration between the twosubsystems increases, such a system may be made increasinglycost-efficient and compact and system calibration will be furthersimplified and more stable over time. Examples for imaging-guided lasersystems in FIGS. 8-16 are integrated at various degrees of integration.

One implementation of a present imaging-guided laser surgical system,for example, includes a surgical laser that produces a surgical laserbeam of surgical laser pulses that cause surgical changes in a targettissue under surgery; a patient interface mount that engages a patientinterface in contact with the target tissue to hold the target tissue inposition; and a laser beam delivery module located between the surgicallaser and the patient interface and configured to direct the surgicallaser beam to the target tissue through the patient interface. Thislaser beam delivery module is operable to scan the surgical laser beamin the target tissue along a predetermined surgical pattern. This systemalso includes a laser control module that controls operation of thesurgical laser and controls the laser beam delivery module to producethe predetermined surgical pattern and an OCT module positioned relativeto the patient interface to have a known spatial relation with respectto the patient interface and the target issue fixed to the patientinterface. The OCT module is configured to direct an optical probe beamto the target tissue and receive returned probe light of the opticalprobe beam from the target tissue to capture OCT images of the targettissue while the surgical laser beam is being directed to the targettissue to perform an surgical operation so that the optical probe beamand the surgical laser beam are simultaneously present in the targettissue. The OCT module is in communication with the laser control moduleto send information of the captured OCT images to the laser controlmodule.

In addition, the laser control module in this particular system respondsto the information of the captured OCT images to operate the laser beamdelivery module in focusing and scanning of the surgical laser beam andadjusts the focusing and scanning of the surgical laser beam in thetarget tissue based on positioning information in the captured OCTimages.

In some implementations, acquiring a complete image of a target tissuemay not be necessary for registering the target to the surgicalinstrument and it may be sufficient to acquire a portion of the targettissue, e.g., a few points from the surgical region such as natural orartificial landmarks. For example, a rigid body has six degrees offreedom in 3D space and six independent points would be sufficient todefine the rigid body. When the exact size of the surgical region is notknown, additional points are needed to provide the positional reference.In this regard, several points can be used to determine the position andthe curvature of the anterior and posterior surfaces, which are normallydifferent, and the thickness and diameter of the crystalline lens of thehuman eye. Based on these data a body made up from two halves ofellipsoid bodies with given parameters can approximate and visualize acrystalline lens for practical purposes. In another implementation,information from the captured image may be combined with informationfrom other sources, such as pre-operative measurements of lens thicknessthat are used as an input for the controller.

FIG. 8 shows one example of an imaging-guided laser surgical system withseparated laser surgical system 2100 and imaging system 2200. The lasersurgical system 2100 includes a laser engine 2130 with a surgical laserthat produces a surgical laser beam 2160 of surgical laser pulses. Alaser beam delivery module 2140 is provided to direct the surgical laserbeam 2160 from the laser engine 2130 to the target tissue 1001 through apatient interface 2150 and is operable to scan the surgical laser beam2160 in the target tissue 1001 along a predetermined surgical pattern. Alaser control module 2120 is provided to control the operation of thesurgical laser in the laser engine 2130 via a communication channel 2121and controls the laser beam delivery module 2140 via a communicationchannel 2122 to produce the predetermined surgical pattern. A patientinterface mount is provided to engage the patient interface 2150 incontact with the target tissue 1001 to hold the target tissue 1001 inposition. The patient interface 2150 can be implemented to include acontact lens or applanation lens with a flat or curved surface toconformingly engage to the anterior surface of the eye and to hold theeye in position.

The imaging system 2200 in FIG. 8 can be an OCT module positionedrelative to the patient interface 2150 of the surgical system 2100 tohave a known spatial relation with respect to the patient interface 2150and the target issue 1001 fixed to the patient interface 2150. This OCTmodule 2200 can be configured to have its own patient interface 2240 forinteracting with the target tissue 1001. The imaging system 2200includes an imaging control module 2220 and an imaging sub-system 2230.The sub-system 2230 includes a light source for generating imaging beam2250 for imaging the target 1001 and an imaging beam delivery module todirect the optical probe beam or imaging beam 2250 to the target tissue1001 and receive returned probe light 2260 of the optical imaging beam2250 from the target tissue 1001 to capture OCT images of the targettissue 1001. Both the optical imaging beam 2250 and the surgical beam2160 can be simultaneously directed to the target tissue 1001 to allowfor sequential or simultaneous imaging and surgical operation.

As illustrated in FIG. 8, communication interfaces 2110 and 2210 areprovided in both the laser surgical system 2100 and the imaging system2200 to facilitate the communications between the laser control by thelaser control module 2120 and imaging by the imaging system 2200 so thatthe OCT module 2200 can send information of the captured OCT images tothe laser control module 2120. The laser control module 2120 in thissystem responds to the information of the captured OCT images to operatethe laser beam delivery module 2140 in focusing and scanning of thesurgical laser beam 2160 and dynamically adjusts the focusing andscanning of the surgical laser beam 2160 in the target tissue 1001 basedon positioning information in the captured OCT images. The integrationbetween the laser surgical system 2100 and the imaging system 2200 ismainly through communication between the communication interfaces 2110and 2210 at the software level.

In this and other examples, various subsystems or devices may also beintegrated. For example, certain diagnostic instruments such aswavefront aberrometers, corneal topography measuring devices may beprovided in the system, or pre-operative information from these devicescan be utilized to augment intra-operative imaging.

FIG. 9 shows an example of an imaging-guided laser surgical system withadditional integration features. The imaging and surgical systems sharea common patient interface 3300 which immobilizes target tissue 1001(e.g., the eye) without having two separate patient interfaces as inFIG. 8. The surgical beam 3210 and the imaging beam 3220 are combined atthe patient interface 3330 and are directed to the target 1001 by thecommon patient interface 3300. In addition, a common control module 3100is provided to control both the imaging sub-system 2230 and the surgicalpart (the laser engine 2130 and the beam delivery system 2140). Thisincreased integration between imaging and surgical parts allows accuratecalibration of the two subsystems and the stability of the position ofthe patient and surgical volume. A common housing 3400 is provided toenclose both the surgical and imaging subsystems. When the two systemsare not integrated into a common housing, the common patient interface3300 can be part of either the imaging or the surgical subsystem.

FIG. 10 shows an example of an imaging-guided laser surgical systemwhere the laser surgical system and the imaging system share both acommon beam delivery module 4100 and a common patient interface 4200.This integration further simplifies the system structure and systemcontrol operation.

In one implementation, the imaging system in the above and otherexamples can be an optical computed tomography (OCT) system and thelaser surgical system is a femtosecond or picosecond laser basedophthalmic surgical system. In OCT, light from a low coherence,broadband light source such as a super luminescent diode is split intoseparate reference and signal beams. The signal beam is the imaging beamsent to the surgical target and the returned light of the imaging beamis collected and recombined coherently with the reference beam to forman interferometer. Scanning the signal beam perpendicularly to theoptical axis of the optical train or the propagation direction of thelight provides spatial resolution in the x-y direction while depthresolution comes from extracting differences between the path lengths ofthe reference arm and the returned signal beam in the signal arm of theinterferometer. While the x-y scanner of different OCT implementationsare essentially the same, comparing the path lengths and getting z-scaninformation can happen in different ways. In one implementation known asthe time domain OCT, for example, the reference arm is continuouslyvaried to change its path length while a photodetector detectsinterference modulation in the intensity of the re-combined beam. In adifferent implementation, the reference arm is essentially static andthe spectrum of the combined light is analyzed for interference. TheFourier transform of the spectrum of the combined beam provides spatialinformation on the scattering from the interior of the sample. Thismethod is known as the spectral domain or Fourier OCT method. In adifferent implementation known as a frequency swept OCT (S. R. Chinn,et. al., Opt. Lett. 22, 1997), a narrowband light source is used withits frequency swept rapidly across a spectral range. Interferencebetween the reference and signal arms is detected by a fast detector anddynamic signal analyzer. An external cavity tuned diode laser orfrequency tuned of frequency domain mode-locked (FDML) laser developedfor this purpose (R. Huber et. Al. Opt. Express, 13, 2005) (S. H. Yun,IEEE J. of Sel. Q. El. 3(4) p. 1087-1096, 1997) can be used in theseexamples as a light source. A femtosecond laser used as a light sourcein an OCT system can have sufficient bandwidth and can provideadditional benefits of increased signal to noise ratios.

The OCT imaging device in the systems in this document can be used toperform various imaging functions. For example, the OCT can be used tosuppress complex conjugates resulting from the optical configuration ofthe system or the presence of the applanation plate, capture OCT imagesof selected locations inside the target tissue to providethree-dimensional positioning information for controlling focusing andscanning of the surgical laser beam inside the target tissue, or captureOCT images of selected locations on the surface of the target tissue oron the applanation plate to provide positioning registration forcontrolling changes in orientation that occur with positional changes ofthe target, such as from upright to supine. The OCT can be calibrated bya positioning registration process based on placement of marks ormarkers in one positional orientation of the target that can then bedetected by the OCT module when the target is in another positionalorientation. In other implementations, the OCT imaging system can beused to produce a probe light beam that is polarized to optically gatherthe information on the internal structure of the eye. The laser beam andthe probe light beam may be polarized in different polarizations. TheOCT can include a polarization control mechanism that controls the probelight used for said optical tomography to polarize in one polarizationwhen traveling toward the eye and in a different polarization whentraveling away from the eye. The polarization control mechanism caninclude, e.g., a wave-plate or a Faraday rotator.

The system in FIG. 10 is shown as a spectral OCT configuration and canbe configured to share the focusing optics part of the beam deliverymodule between the surgical and the imaging systems. The mainrequirements for the optics are related to the operating wavelength,image quality, resolution, distortion etc. The laser surgical system canbe a femtosecond laser system with a high numerical aperture systemdesigned to achieve diffraction limited focal spot sizes, e.g., about 2to 3 micrometers. Various femtosecond ophthalmic surgical lasers canoperate at various wavelengths such as wavelengths of around 1.05micrometer. The operating wavelength of the imaging device can beselected to be close to the laser wavelength so that the optics ischromatically compensated for both wavelengths. Such a system mayinclude a third optical channel, a visual observation channel such as asurgical microscope, to provide an additional imaging device to captureimages of the target tissue. If the optical path for this third opticalchannel shares optics with the surgical laser beam and the light of theOCT imaging device, the shared optics can be configured with chromaticcompensation in the visible spectral band for the third optical channeland the spectral bands for the surgical laser beam and the OCT imagingbeam.

FIG. 11 shows a particular example of the design in FIG. 9 where thescanner 5100 for scanning the surgical laser beam and the beamconditioner 5200 for conditioning (collimating and focusing) thesurgical laser beam are separate from the optics in the OCT imagingmodule 5300 for controlling the imaging beam for the OCT. The surgicaland imaging systems share an objective lens 5600 module and the patientinterface 3300. The objective lens 5600 directs and focuses both thesurgical laser beam and the imaging beam to the patient interface 3300and its focusing is controlled by the control module 3100. Two beamsplitters 5410 and 5420 are provided to direct the surgical and imagingbeams. The beam splitter 5420 is also used to direct the returnedimaging beam back into the OCT imaging module 5300. Two beam splitters5410 and 5420 also direct light from the target 1001 to a visualobservation optics unit 5500 to provide direct view or image of thetarget 1001. The unit 5500 can be a lens imaging system for the surgeonto view the target 1001 or a camera to capture the image or video of thetarget 1001. Various beam splitters can be used, such as dichroic andpolarization beam splitters, optical grating, holographic beam splitteror a combinations of these.

In some implementations, the optical components may be appropriatelycoated with antireflection coating for both the surgical and for the OCTwavelength to reduce glare from multiple surfaces of the optical beampath. Reflections would otherwise reduce the throughput of the systemand reduce the signal to noise ratio by increasing background light inthe OCT imaging unit. One way to reduce glare in the OCT is to rotatethe polarization of the return light from the sample by wave-plate ofFaraday isolator placed close to the target tissue and orient apolarizer in front of the OCT detector to preferentially detect lightreturned from the sample and suppress light scattered from the opticalcomponents.

In a laser surgical system, each of the surgical laser and the OCTsystem can have a beam scanner to cover the same surgical region in thetarget tissue. Hence, the beam scanning for the surgical laser beam andthe beam scanning for the imaging beam can be integrated to share commonscanning devices.

FIG. 12 shows an example of such a system in detail. In thisimplementation the x-y scanner 6410 and the z scanner 6420 are shared byboth subsystems. A common control 6100 is provided to control the systemoperations for both surgical and imaging operations. The OCT sub-systemincludes an OCT light source 6200 that produce the imaging light that issplit into an imaging beam and a reference beam by a beam splitter 6210.The imaging beam is combined with the surgical beam at the beam splitter6310 to propagate along a common optical path leading to the target1001. The scanners 6410 and 6420 and the beam conditioner unit 6430 arelocated downstream from the beam splitter 6310. A beam splitter 6440 isused to direct the imaging and surgical beams to the objective lens 5600and the patient interface 3300.

In the OCT sub-system, the reference beam transmits through the beamsplitter 6210 to an optical delay device 6220 and is reflected by areturn mirror 6230. The returned imaging beam from the target 1001 isdirected back to the beam splitter 6310 which reflects at least aportion of the returned imaging beam to the beam splitter 6210 where thereflected reference beam and the returned imaging beam overlap andinterfere with each other. A spectrometer detector 6240 is used todetect the interference and to produce OCT images of the target 1001.The OCT image information is sent to the control system 6100 forcontrolling the surgical laser engine 2130, the scanners 6410 and 6420and the objective lens 5600 to control the surgical laser beam. In oneimplementation, the optical delay device 6220 can be varied to changethe optical delay to detect various depths in the target tissue 1001.

If the OCT system is a time domain system, the two subsystems use twodifferent z-scanners because the two scanners operate in different ways.In this example, the z scanner of the surgical system operates bychanging the divergence of the surgical beam in the beam conditionerunit without changing the path lengths of the beam in the surgical beampath. On the other hand, the time domain OCT scans the z-direction byphysically changing the beam path by a variable delay or by moving theposition of the reference beam return mirror. After calibration, the twoz-scanners can be synchronized by the laser control module. Therelationship between the two movements can be simplified to a linear orpolynomial dependence, which the control module can handle oralternatively calibration points can define a look-up table to provideproper scaling. Spectral/Fourier domain and frequency swept source OCTdevices have no z-scanner, the length of the reference arm is static.Besides reducing costs, cross calibration of the two systems will berelatively straightforward. There is no need to compensate fordifferences arising from image distortions in the focusing optics orfrom the differences of the scanners of the two systems since they areshared.

In practical implementations of the surgical systems, the focusingobjective lens 5600 is slidably or movably mounted on a base and theweight of the objective lens is balanced to limit the force on thepatient's eye. The patient interface 3300 can include an applanationlens attached to a patient interface mount. The patient interface mountis attached to a mounting unit, which holds the focusing objective lens.This mounting unit is designed to ensure a stable connection between thepatient interface and the system in case of unavoidable movement of thepatient and allows gentler docking of the patient interface onto theeye. Various implementations for the focusing objective lens can be usedand one example is described in U.S. Pat. No. 5,336,215 to Hsueh. Thispresence of an adjustable focusing objective lens can change the opticalpath length of the optical probe light as part of the opticalinterferometer for the OCT sub-system. Movement of the objective lens5600 and patient interface 3300 can change the path length differencesbetween the reference beam and the imaging signal beam of the OCT in anuncontrolled way and this may degrade the OCT depth information detectedby the OCT. This would happen not only in time-domain but also inspectral/Fourier domain and frequency-swept OCT systems.

FIGS. 13 and 14 show exemplary imaging-guided laser surgical systemsthat address the technical issue associated with the adjustable focusingobjective lens.

The system in FIG. 13 provides a position sensing device 7110 coupled tothe movable focusing objective lens 7100 to measure the position of theobjective lens 7100 on a slideable mount and communicates the measuredposition to a control module 7200 in the OCT system. The control system6100 can control and move the position of the objective lens 7100 toadjust the optical path length traveled by the imaging signal beam forthe OCT operation and the position of the lens 7100 is measured andmonitored by the position encoder 7110 and direct fed to the OCT control7200. The control module 7200 in the OCT system applies an algorithm,when assembling a 3D image in processing the OCT data, to compensate fordifferences between the reference arm and the signal arm of theinterferometer inside the OCT caused by the movement of the focusingobjective lens 7100 relative to the patient interface 3300. The properamount of the change in the position of the lens 7100 computed by theOCT control module 7200 is sent to the control 6100 which controls thelens 7100 to change its position.

FIG. 14 shows another exemplary system where the return mirror 6230 inthe reference arm of the interferometer of the OCT system or at leastone part in an optical path length delay assembly of the OCT system isrigidly attached to the movable focusing objective lens 7100 so thesignal arm and the reference arm undergo the same amount of change inthe optical path length when the objective lens 7100 moves. As such, themovement of the objective lens 7100 on the slide is automaticallycompensated for path-length differences in the OCT system withoutadditional need for a computational compensation.

The above examples for imaging-guided laser surgical systems, the lasersurgical system and the OCT system use different light sources. In aneven more complete integration between the laser surgical system and theOCT system, a femtosecond surgical laser as a light source for thesurgical laser beam can also be used as the light source for the OCTsystem.

FIG. 15 shows an example where a femtosecond pulse laser in a lightmodule 9100 is used to generate both the surgical laser beam forsurgical operations and the probe light beam for OCT imaging. A beamsplitter 9300 is provided to split the laser beam into a first beam asboth the surgical laser beam and the signal beam for the OCT and asecond beam as the reference beam for the OCT. The first beam isdirected through an x-y scanner 6410 which scans the beam in the x and ydirections perpendicular to the propagation direction of the first beamand a second scanner (z scanner) 6420 that changes the divergence of thebeam to adjust the focusing of the first beam at the target tissue 1001.This first beam performs the surgical operations at the target tissue1001 and a portion of this first beam is back scattered to the patientinterface and is collected by the objective lens as the signal beam forthe signal arm of the optical interferometer of the OCT system. Thisreturned light is combined with the second beam that is reflected by areturn mirror 6230 in the reference arm and is delayed by an adjustableoptical delay element 6220 for a time-domain OCT to control the pathdifference between the signal and reference beams in imaging differentdepths of the target tissue 1001. The control system 9200 controls thesystem operations.

Surgical practice on the cornea has shown that a pulse duration ofseveral hundred femtoseconds may be sufficient to achieve good surgicalperformance, while for OCT of a sufficient depth resolution broaderspectral bandwidth generated by shorter pulses, e.g., below several tensof femtoseconds, are needed. In this context, the design of the OCTdevice dictates the duration of the pulses from the femtosecond surgicallaser.

FIG. 16 shows another imaging-guided system that uses a single pulsedlaser 9100 to produce the surgical light and the imaging light. Anonlinear spectral broadening media 9400 is placed in the output opticalpath of the femtosecond pulsed laser to use an optical non-linearprocess such as white light generation or spectral broadening to broadenthe spectral bandwidth of the pulses from a laser source of relativelylonger pulses, several hundred femtoseconds normally used in surgery.The media 9400 can be a fiber-optic material, for example. The lightintensity requirements of the two systems are different and a mechanismto adjust beam intensities can be implemented to meet such requirementsin the two systems. For example, beam steering mirrors, beam shutters orattenuators can be provided in the optical paths of the two systems toproperly control the presence and intensity of the beam when taking anOCT image or performing surgery in order to protect the patient andsensitive instruments from excessive light intensity.

In operation, the above examples in FIGS. 8/16 can be used to performimaging-guided laser surgery. FIG. 17 shows one example of a method forperforming laser surgery by using an imaging-guided laser surgicalsystem. This method uses a patient interface in the system to engage toand to hold a target tissue under surgery in position and simultaneouslydirects a surgical laser beam of laser pulses from a laser in the systemand an optical probe beam from the OCT module in the system to thepatient interface into the target tissue. The surgical laser beam iscontrolled to perform laser surgery in the target tissue and the OCTmodule is operated to obtain OCT images inside the target tissue fromlight of the optical probe beam returning from the target tissue. Theposition information in the obtained OCT images is applied in focusingand scanning of the surgical laser beam to adjust the focusing andscanning of the surgical laser beam in the target tissue before orduring surgery.

FIG. 18 shows an example of an OCT image of an eye. The contactingsurface of the applanation lens in the patent interface can beconfigured to have a curvature that minimizes distortions or folds inthe cornea due to the pressure exerted on the eye during applanation.After the eye is successfully applanated at the patient interface, anOCT image can be obtained. As illustrated in FIG. 18, the curvature ofthe lens and cornea as well as the distances between the lens and corneaare identifiable in the OCT image. Subtler features such as theepithelium-cornea interface are detectable. Each of these identifiablefeatures may be used as an internal reference of the laser coordinateswith the eye. The coordinates of the cornea and lens can be digitizedusing well-established computer vision algorithms such as Edge or Blobdetection. Once the coordinates of the lens are established, they can beused to control the focusing and positioning of the surgical laser beamfor the surgery.

Alternatively, a calibration sample material may be used to form a 3-Darray of reference marks at locations with known position coordinates.The OCT image of the calibration sample material can be obtained toestablish a mapping relationship between the known position coordinatesof the reference marks and the OCT images of the reference marks in theobtained OCT image. This mapping relationship is stored as digitalcalibration data and is applied in controlling the focusing and scanningof the surgical laser beam during the surgery in the target tissue basedon the OCT images of the target tissue obtained during the surgery. TheOCT imaging system is used here as an example and this calibration canbe applied to images obtained via other imaging techniques.

In an imaging-guided laser surgical system described here, the surgicallaser can produce relatively high peak powers sufficient to drive strongfield/multi-photon ionization inside of the eye (i.e. inside of thecornea and lens) under high numerical aperture focusing. Under theseconditions, one pulse from the surgical laser generates a plasma withinthe focal volume. Cooling of the plasma results in a well defined damagezone or “bubble” that may be used as a reference point. The followingsections describe a calibration procedure for calibrating the surgicallaser against an OCT-based imaging system using the damage zones createdby the surgical laser.

Before surgery can be performed, the OCT is calibrated against thesurgical laser to establish a relative positioning relationship so thatthe surgical laser can be controlled in position at the target tissuewith respect to the position associated with images in the OCT image ofthe target tissue obtained by the OCT. One way for performing thiscalibration uses a pre-calibrated target or “phantom” which can bedamaged by the laser as well as imaged with the OCT. The phantom can befabricated from various materials such as a glass or hard plastic (e.g.PMMA) such that the material can permanently record optical damagecreated by the surgical laser. The phantom can also be selected to haveoptical or other properties (such as water content) that are similar tothe surgical target.

The phantom can be, e.g., a cylindrical material having a diameter of atleast 10 mm (or that of the scanning range of the delivery system) and acylindrical length of at least 10 mm long spanning the distance of theepithelium to the crystalline lens of the eye, or as long as thescanning depth of the surgical system. The upper surface of the phantomcan be curved to mate seamlessly with the patient interface or thephantom material may be compressible to allow full applanation. Thephantom may have a three dimensional grid such that both the laserposition (in x and y) and focus (z), as well as the OCT image can bereferenced against the phantom.

FIGS. 19A-19D illustrate two exemplary configurations for the phantom.FIG. 19A illustrates a phantom that is segmented into thin disks. FIG.19B shows a single disk patterned to have a grid of reference marks as areference for determining the laser position across the phantom (i.e.the x- and y-coordinates). The z-coordinate (depth) can be determined byremoving an individual disk from the stack and imaging it under aconfocal microscope.

FIG. 19C illustrates a phantom that can be separated into two halves.Similar to the segmented phantom in FIG. 19A, this phantom is structuredto contain a grid of reference marks as a reference for determining thelaser position in the x- and y-coordinates. Depth information can beextracted by separating the phantom into the two halves and measuringthe distance between damage zones. The combined information can providethe parameters for image guided surgery.

FIG. 20 shows a surgical system part of the imaging-guided lasersurgical system. This system includes steering mirrors which may beactuated by actuators such as galvanometers or voice coils, an objectivelens e and a disposable patient interface. The surgical laser beam isreflected from the steering mirrors through the objective lens. Theobjective lens focuses the beam just after the patient interface.Scanning in the x- and y-coordinates is performed by changing the angleof the beam relative to the objective lens. Scanning in z-plane isaccomplished by changing the divergence of the incoming beam using asystem of lens upstream to the steering mirrors.

In this example, the conical section of the disposable patient interfacemay be either air spaced or solid and the section interfacing with thepatient includes a curved contact lens. The curved contact lens can befabricated from fused silica or other material resistant to formingcolor centers when irradiated with ionizing radiation. The radius ofcurvature is on the upper limit of what is compatible with the eye,e.g., about 10 mm.

The first step in the calibration procedure is docking the patientinterface with the phantom. The curvature of the phantom matches thecurvature of the patient interface. After docking, the next step in theprocedure involves creating optical damage inside of the phantom toproduce the reference marks.

FIG. 21 shows examples of actual damage zones produced by a femtosecondlaser in glass. The separation between the damage zones is on average 8μm (the pulse energy is 2.2 μJ with duration of 580 fs at full width athalf maximum). The optical damage depicted in FIG. 21 shows that thedamage zones created by the femtosecond laser are well-defined anddiscrete. In the example shown, the damage zones have a diameter ofabout 2.5 μm. Optical damage zones similar to that shown in FIG. 20 arecreated in the phantom at various depths to form a 3-D array of thereference marks. These damage zones are referenced against thecalibrated phantom either by extracting the appropriate disks andimaging it under a confocal microscope (FIG. 19A) or by splitting thephantom into two halves and measuring the depth using a micrometer (FIG.19C). The x- and y-coordinates can be established from thepre-calibrated grid.

After damaging the phantom with the surgical laser, OCT on the phantomis performed. The OCT imaging system provides a 3D rendering of thephantom establishing a relationship between the OCT coordinate systemand the phantom. The damage zones are detectable with the imagingsystem. The OCT and laser may be cross-calibrated using the phantom'sinternal standard. After the OCT and the laser are referenced againsteach other, the phantom can be discarded.

Prior to surgery, the calibration can be verified. This verificationstep involves creating optical damage at various positions inside of asecond phantom. The optical damage should be intense enough such thatthe multiple damage zones which create a circular pattern can be imagedby the OCT. After the pattern is created, the second phantom is imagedwith the OCT. Comparison of the OCT image with the laser coordinatesprovides the final check of the system calibration prior to surgery.

Once the coordinates are fed into the laser, laser surgery can beperformed inside the eye. This involves photo-emulsification of the lensusing the laser, as well as other laser treatments to the eye. Thesurgery can be stopped at any time and the anterior segment of the eye(FIG. 17) can be re-imaged to monitor the progress of the surgery;moreover, after the IOL is inserted, imaging the IOL (with light or noapplanation) provides information regarding the position of the IOL inthe eye. This information may be utilized by the physician to refine theposition of the IOL.

FIG. 22 shows an example of the calibration process and thepost-calibration surgical operation. This examples illustrates a methodfor performing laser surgery by using an imaging-guided laser surgicalsystem can include using a patient interface in the system, that isengaged to hold a target tissue under surgery in position, to hold acalibration sample material during a calibration process beforeperforming a surgery; directing a surgical laser beam of laser pulsesfrom a laser in the system to the patient interface into the calibrationsample material to burn reference marks at selected three-dimensionalreference locations; directing an optical probe beam from an opticalcoherence tomography (OCT) module in the system to the patient interfaceinto the calibration sample material to capture OCT images of the burntreference marks; and establishing a relationship between positioningcoordinates of the OCT module and the burnt reference marks. After theestablishing the relationship, a patient interface in the system is usedto engage to and to hold a target tissue under surgery in position. Thesurgical laser beam of laser pulses and the optical probe beam aredirected to the patient interface into the target tissue. The surgicallaser beam is controlled to perform laser surgery in the target tissue.The OCT module is operated to obtain OCT images inside the target tissuefrom light of the optical probe beam returning from the target tissueand the position information in the obtained OCT images and theestablished relationship are applied in focusing and scanning of thesurgical laser beam to adjust the focusing and scanning of the surgicallaser beam in the target tissue during surgery. While such calibrationscan be performed immediately prior to laser surgery, they can also beperformed at various intervals before a procedure, using calibrationvalidations that demonstrated a lack of drift or change in calibrationduring such intervals.

The following examples describe imaging-guided laser surgical techniquesand systems that use images of laser-induced photodisruption byproductsfor alignment of the surgical laser beam.

FIGS. 23A and 23B illustrate another implementation of the presenttechnique in which actual photodisruption byproducts in the targettissue are used to guide further laser placement. A pulsed laser 1710,such as a femtosecond or picosecond laser, is used to produce a laserbeam 1712 with laser pulses to cause photodisruption in a target tissue1001. The target tissue 1001 may be a part of a body part 1700 of asubject, e.g., a portion of the lens of one eye. The laser beam 1712 isfocused and directed by an optics module for the laser 1710 to a targettissue position in the target tissue 1001 to achieve a certain surgicaleffect. The target surface is optically coupled to the laser opticsmodule by an applanation plate 1730 that transmits the laser wavelength,as well as image wavelengths from the target tissue. The applanationplate 1730 can be an applanation lens. An imaging device 1720 isprovided to collect reflected or scattered light or sound from thetarget tissue 1001 to capture images of the target tissue 1001 eitherbefore or after (or both) the applanation plate is applied. The capturedimaging data is then processed by the laser system control module todetermine the desired target tissue position. The laser system controlmodule moves or adjusts optical or laser elements based on standardoptical models to ensure that the center of photodisruption byproduct1702 overlaps with the target tissue position. This can be a dynamicalignment process where the images of the photodisruption byproduct 1702and the target tissue 1001 are continuously monitored during thesurgical process to ensure that the laser beam is properly positioned ateach target tissue position.

In one implementation, the laser system can be operated in two modes:first in a diagnostic mode in which the laser beam 1712 is initiallyaligned by using alignment laser pulses to create photodisruptionbyproduct 1702 for alignment and then in a surgical mode where surgicallaser pulses are generated to perform the actual surgical operation. Inboth modes, the images of the disruption byproduct 1702 and the targettissue 1001 are monitored to control the beam alignment. FIG. 17A showsthe diagnostic mode where the alignment laser pulses in the laser beam1712 may be set at a different energy level than the energy level of thesurgical laser pulses. For example, the alignment laser pulses may beless energetic than the surgical laser pulses but sufficient to causesignificant photodisruption in the tissue to capture the photodisruptionbyproduct 1702 at the imaging device 1720. The resolution of this coarsetargeting may not be sufficient to provide desired surgical effect.Based on the captured images, the laser beam 1712 can be alignedproperly. After this initial alignment, the laser 1710 can be controlledto produce the surgical laser pulses at a higher energy level to performthe surgery. Because the surgical laser pulses are at a different energylevel than the alignment laser pulses, the nonlinear effects in thetissue material in the photodisruption can cause the laser beam 1712 tobe focused at a different position from the beam position during thediagnostic mode. Therefore, the alignment achieved during the diagnosticmode is a coarse alignment and additional alignment can be furtherperformed to precisely position each surgical laser pulse during thesurgical mode when the surgical laser pulses perform the actual surgery.Referring to FIG. 23A, the imaging device 1720 captures the images fromthe target tissue 1001 during the surgical mode and the laser controlmodule adjust the laser beam 1712 to place the focus position 1714 ofthe laser beam 1712 onto the desired target tissue position in thetarget tissue 1001. This process is performed for each target tissueposition.

FIG. 24 shows one implementation of the laser alignment where the laserbeam is first approximately aimed at the target tissue and then theimage of the photodisruption byproduct is captured and used to align thelaser beam. The image of the target tissue of the body part as thetarget tissue and the image of a reference on the body part aremonitored to aim the pulsed laser beam at the target tissue. The imagesof photodisruption byproduct and the target tissue are used to adjustthe pulsed laser beam to overlap the location of the photodisruptionbyproduct with the target tissue.

FIG. 25 shows one implementation of the laser alignment method based onimaging photodisruption byproduct in the target tissue in laser surgery.In this method, a pulsed laser beam is aimed at a target tissue locationwithin target tissue to deliver a sequence of initial alignment laserpulses to the target tissue location. The images of the target tissuelocation and photodisruption byproduct caused by the initial alignmentlaser pulses are monitored to obtain a location of the photodisruptionbyproduct relative to the target tissue location. The location ofphotodisruption byproduct caused by surgical laser pulses at a surgicalpulse energy level different from the initial alignment laser pulses isdetermined when the pulsed laser beam of the surgical laser pulses isplaced at the target tissue location. The pulsed laser beam iscontrolled to carry surgical laser pulses at the surgical pulse energylevel. The position of the pulsed laser beam is adjusted at the surgicalpulse energy level to place the location of photodisruption byproduct atthe determined location. While monitoring images of the target tissueand the photodisruption byproduct, the position of the pulsed laser beamat the surgical pulse energy level is adjusted to place the location ofphotodisruption byproduct at a respective determined location whenmoving the pulsed laser beam to a new target tissue location within thetarget tissue.

FIG. 26 shows an exemplary laser surgical system based on the laseralignment using the image of the photodisruption byproduct. An opticsmodule 2010 is provided to focus and direct the laser beam to the targettissue 1700. The optics module 2010 can include one or more lenses andmay further include one or more reflectors. A control actuator isincluded in the optics module 2010 to adjust the focusing and the beamdirection in response to a beam control signal. A system control module2020 is provided to control both the pulsed laser 1010 via a lasercontrol signal and the optics module 2010 via the beam control signal.The system control module 2020 processes image data from the imagingdevice 2030 that includes the position offset information for thephotodisruption byproduct 1702 from the target tissue position in thetarget tissue 1700. Based on the information obtained from the image,the beam control signal is generated to control the optics module 2010which adjusts the laser beam. A digital processing unit is included inthe system control module 2020 to perform various data processing forthe laser alignment.

The imaging device 2030 can be implemented in various forms, includingan optical coherent tomography (OCT) device. In addition, an ultrasoundimaging device can also be used. The position of the laser focus ismoved so as to place it grossly located at the target at the resolutionof the imaging device. The error in the referencing of the laser focusto the target and possible non-linear optical effects such as selffocusing that make it difficult to accurately predict the location ofthe laser focus and subsequent photodisruption event. Variouscalibration methods, including the use of a model system or softwareprogram to predict focusing of the laser inside a material can be usedto get a coarse targeting of the laser within the imaged tissue. Theimaging of the target can be performed both before and after thephotodisruption. The position of the photodisruption by productsrelative to the target is used to shift the focal point of the laser tobetter localize the laser focus and photodisruption process at orrelative to the target. Thus the actual photodisruption event is used toprovide a precise targeting for the placement of subsequent surgicalpulses.

Photodisruption for targeting during the diagnostic mode can beperformed at a lower, higher or the same energy level that is requiredfor the later surgical processing in the surgical mode of the system. Acalibration may be used to correlate the localization of thephotodisruptive event performed at a different energy in diagnostic modewith the predicted localization at the surgical energy because theoptical pulse energy level can affect the exact location of thephotodisruptive event. Once this initial localization and alignment isperformed, a volume or pattern of laser pulses (or a single pulse) canbe delivered relative to this positioning. Additional sampling imagescan be made during the course of delivering the additional laser pulsesto ensure proper localization of the laser (the sampling images may beobtained with use of lower, higher or the same energy pulses). In oneimplementation, an ultrasound device is used to detect the cavitationbubble or shock wave or other photodisruption byproduct. Thelocalization of this can then be correlated with imaging of the target,obtained via ultrasound or other modality. In another embodiment, theimaging device is simply a biomicroscope or other optical visualizationof the photodisruption event by the operator, such as optical coherencetomography. With the initial observation, the laser focus is moved tothe desired target position, after which a pattern or volume of pulsesis delivered relative to this initial position.

As a specific example, a laser system for precise subsurfacephotodisruption can include means for generating laser pulses capable ofgenerating photodisruption at repetition rates of 100-1000 Millionpulses per second, means for coarsely focusing laser pulses to a targetbelow a surface using an image of the target and a calibration of thelaser focus to that image without creating a surgical effect, means fordetecting or visualizing below a surface to provide an image orvisualization of a target the adjacent space or material around thetarget and the byproducts of at least one photodisruptive event coarselylocalized near the target, means for correlating the position of thebyproducts of photodisruption with that of the sub surface target atleast once and moving the focus of the laser pulse to position thebyproducts of photodisruption at the sub surface target or at a relativeposition relative to the target, means for delivering a subsequent trainof at least one additional laser pulse in pattern relative to theposition indicated by the above fine correlation of the byproducts ofphotodisruption with that of the sub surface target, and means forcontinuing to monitor the photodisruptive events during placement of thesubsequent train of pulses to further fine tune the position of thesubsequent laser pulses relative to the same or revised target beingimaged.

The above techniques and systems can be used deliver high repetitionrate laser pulses to subsurface targets with a precision required forcontiguous pulse placement, as needed for cutting or volume disruptionapplications. This can be accomplished with or without the use of areference source on the surface of the target and can take into accountmovement of the target following applanation or during placement oflaser pulses.

While this specification described various embodiments andimplementations, these should not be construed as limitations on thescope of an invention or of what may be claimed, but rather asdescriptions of features specific to particular embodiments of theinvention. Certain features that are described in this specification inthe context of separate embodiments can also be implemented incombination in a single embodiment. Conversely, various features thatare described in the context of a single embodiment can also beimplemented in multiple embodiments separately or in any suitablesub-combination. Moreover, although features may be described above asacting in certain combinations and even initially claimed as such, oneor more features from a claimed combination can in some cases be excisedfrom the combination, and the claimed combination may be directed to asub-combination or a variation of a sub-combination. Also, enhancements,combinations, extensions and variations can be made based on what isdisclosed and illustrated.

What is claimed is:
 1. A laser system for fragmenting a crystalline lensof an eye, comprising: a pulsed laser that generates a laser beam oflaser pulses; and an optical delivery system comprising: optics with acurved focal plane that receive the laser beam and focus the laser beamonto the crystalline lens, the optics comprising one or more lenses; anXY scanner that scans the laser beam in x and y directions perpendicularto the propagation direction of the laser beam with a first scanningspeed; and a Z-scanner that scans the laser beam in a z directionparallel to the propagation direction of the laser beam with a secondscanning speed slower than the first scanning speed, wherein the opticaldelivery system applies the laser beam to create incisions forming aplurality of layers in the crystalline lens of the eye on alayer-by-layer basis by: scanning the laser beam with the XY scanneralong the curved focal plane of the optics to create an incision of alayer that follows the curvature of the curved focal plane; and scanningthe laser beam with the Z scanner at the second scanning speed slowerthan the first scanning speed of the XY scanner to create the incisionof the layer, without adjusting the Z scanner at the first scanningspeed of the XY scanner.
 2. The laser system of claim 1, wherein theoptical delivery system is configured to move a focal point of the laserin a posterior-to-anterior direction of the lens.
 3. The laser system ofclaim 1, wherein the optical delivery system is configured to controlthe laser to generate a laser beam with laser-parameters: sufficient tocreate photodisruption in a selected lens region; and insufficient tocause damage to a retina of the eye.
 4. The laser system of claim 3,wherein the optical delivery system is configured to control the pulsedlaser to generate laser pulses with laser-parameters comprising: anenergy in the range of 0.5 microJ to 50 microJ; a separation of adjacenttarget areas in the range of 1 micron to 100 microns; a duration in therange of 0.005 picoseconds to 25 picoseconds; and a repetition rate inthe range of 1 kHz to 10 MHz.
 5. The laser system of claim 1, wherein:the incisions are formed on a layer-by-layer basis without interruptingthe application of the laser.
 6. The system of claim 1, wherein anorientation of a portion of the incision is one of an orientationintersecting fibers of the lens and an orientation non-transverse to anaxis of the eye.
 7. The system of claim 6, wherein the incision has aspatial extent in a Z direction in the range of 0.5-10 mm, and in an X-Yplane in the range of 2-10 mm.
 8. The system of claim 6, wherein thenon-transverse orientation of the incision is one of: an orientationsubstantially parallel to the axis of the eye; and an orientation makinga less than 90 degree angle with the axis of the eye.
 9. The lasersystem of claim 1, wherein forming the incisions on a layer-by-layerbasis comprises: applying laser pulses to target locations within aposterior layer of the lens, the target locations belonging to twoincisions or two segments of the same incision; and applying laserpulses to target locations within a layer anterior to the posteriorlayer, the target locations belonging to the same two incisions or tothe same two segments of the same incision.
 10. The system of claim 1,wherein the non-transverse orientation of the incision is one of: anorientation substantially parallel to the axis of the eye; and anorientation making a less than 90 degree angle with the axis of the eye.11. The system of claim 1, wherein an orientation of the incision isaligned with a preferential direction of expansion of the bubbles. 12.The system of claim 1, wherein the laser pulses are scanned in acontinuous manner to form the incision without repositioning the laseror interrupting the application of the laser.
 13. The system of claim 1,wherein the incision has a form aligned with the axis of the eye, theform being of at least one of: a cylinder, a set of concentriccylinders, a set of cylinders connected by one or more connecting line,a curved surface, a cone, a spiral, a layered spiral with smooth linesconnecting layers of the spiral and a tilted cylinder.
 14. The system ofclaim 1, wherein the incision has a form aligned with the axis of theeye, the form being at least one of: a plane, two or more crossingplanes, a combination of planes and connecting arcs, and a combinationof planes and cylinders.
 15. The system of claim 1, wherein the opticaldelivery system is configured to scan the laser beam by: applying thelaser pulses to form a first ring with a first radius in a posteriorlayer of the lens; applying the laser pulses to form a connector linebetween the first and a second ring in the posterior layer; applying thelaser pulses to form the second ring with a second radius in theposterior layer; and repeating multiple times the formation of the firstring, the second ring and the connector line in layers sequentiallyanterior to the posterior layer, wherein the first rings in thesequential layers form a first cylinder, the second rings form a secondcylinder, the cylinders being connected by the connector lines.
 16. Thesystem of claim 15, wherein the connector lines in sequential layers areone of: aligned to form connector planes; and not-aligned from layer tolayer.
 17. The system claim 1, wherein the optical delivery system isconfigured to scan the laser beam to: form a posterior spiral in aposterior layer; form a smooth connector line starting near an end ofthe posterior spiral in the posterior layer, the connector line smoothlybending and rising to a central region of a layer anterior to theposterior layer; and form an anterior spiral starting at the end of thesmooth connector line in the central region of the anterior layer. 18.The system of claim 17, wherein the posterior spiral and the anteriorspiral are essentially aligned to form a spiral with an extent in the Zdirection.